Imaging probe with combined ultrasound and optical means of imaging

ABSTRACT

The present invention provides an imaging probe for imaging mammalian tissues and structures using high resolution imaging, including high frequency ultrasound and optical coherence tomography. The imaging probes structures using high resolution imaging use combined high frequency ultrasound (IVUS) and optical imaging methods such as optical coherence tomography (OCT) and to accurate co-registering of images obtained from ultrasound image signals and optical image signals during scanning a region of interest.

CROSS REFERENCE TO RELATED U.S. APPLICATIONS

This patent application relates to, and claims the priority benefitfrom, U.S. patent application Ser. No. 12/010,208 filed on Jan. 22,2008, in English, entitled IMAGING PROBE WITH COMBINED ULTRASOUND ANDOPTICAL MEANS OF IMAGING, which relates to, and claims the prioritybenefit from U.S. Provisional Patent Application Ser. No. 60/881,169filed on Jan. 19, 2007, in English, entitled IMAGING PROBE, the contentsof these applications being incorporated herein by reference in theirentirety.

FIELD OF THE INVENTION

The present invention relates generally to the field of imagingmammalian tissues and structures using high resolution imaging usingcombined high frequency ultrasound (IVUS) and optical imaging methodssuch as optical coherence tomography (OCT) and to accurateco-registering of images obtained from ultrasound image signals andoptical image signals during scanning a region of interest.

BACKGROUND OF THE INVENTION

High resolution imaging of the interior of the body (or for dermatologicor ophthalmology applications not restricted to the interior) servesmultiple purposes, including any of i) assessing tissue structures,anatomy and composition; ii) planning and/or guiding interventions onlocalized regions of the body; and iii) assessing the result ofinterventions that alter the structure, composition or other propertiesof the localized region. High resolution imaging in this particular caserefers to high frequency ultrasound and optical imaging methods. For thepurposes of this invention, high frequency ultrasound typically refersto imaging with frequencies of greater than 3 MHz, and more typically inthe range of 9 to 100 MHz.

High frequency ultrasound is very useful for intravascular andintracardiac procedures. For these applications, the ultrasoundtransducers are incorporated into a catheter or other device that can beinserted into the body. By way of example, two particularly importantimplementations of high frequency ultrasound are intravascularultrasound (IVUS), for imaging blood vessels, and intracardiacechocardiography (ICE) for imaging cardiac chambers. Both ICE and IVUSare minimally invasive, and involve placing one or more ultrasoundtransducers inside a blood vessel or cardiac chamber to take highquality images of these structures.

Optical imaging methods based on fiber optic technology used in thefield of medicine include optical coherence tomography (OCT),angioscopy, near infrared spectroscopy, Raman spectroscopy andfluorescence spectroscopy. These modalities typically require the use ofone or more optical fibers to transmit light energy along a shaftbetween an imaging site and an imaging detector. Optical coherencetomography is an optical analog of ultrasound, and provides imagingresolutions on the order of 1 to 30 microns, but does not penetrate asdeeply into tissue as ultrasound in most cases. Fiber optics can also beused to deliver energy for therapeutic maneuvers such as laser ablationof tissue and photodynamic therapy.

Additional forms of imaging related to this invention includeangioscopy, endoscopy and other similar imaging mechanisms that involvesimaging a site inside the patient using a probe to take pictures basedon the back-reflection of light.

High resolution imaging means have been implemented in many forms forassessing several different regions of mammalian anatomy, including thegastrointestinal system, the cardiovascular system (including coronary,peripheral and neurological vasculature), skin, eyes (including theretina), the genitourinary systems, breast tissue, liver tissue and manyothers. By way of example, imaging of the cardiovascular system withhigh frequency ultrasound or optical coherence tomography has beendeveloped for assessing the structure and composition of arterialplaque.

High-resolution imaging has been used to measure vessel or plaquegeometry, blood flow through diseased arteries, the effects ofinterventions on arterial plaque (such as by atherectomy, angioplastyand/or stenting). Attempts have also been made using high resolutionimaging to identify vascular lesions that have not led to clinicalsymptoms, but are at increased risk of rupturing or eroding and causingan acute myocardial infarction. These so-called “vulnerable plaques” arean area of interest as the prospect of treating such plaques to pre-emptadverse clinical events is conceptually appealing.

Chronic total occlusions are a specific subset of vascular lesions wherethe entire lumen of the vessel has been occluded (based on theangiographic appearance of the lesion) for over approximately one month.Most intravascular imaging modalities are “side-viewing” and requirepassage of an intravascular imaging device through a lesion. In order toimage chronic total occlusions, methods of high resolution imaging wouldbe more useful if they were adapted to a “forward-looking” rather than“side-viewing” configuration.

Several of these high resolution imaging means are dependent on the useof a rotary shaft to transmit torque to an imaging device near thedistal end of the probe. These rotary shafts are often long, thin andflexible, such that they can be delivered through anatomical conduits,such as the vasculature, genitourinary tracts, respiratory tracts andother such bodily lumens. Ideally, when a continuous torque is appliedto the cable in a specified direction the torque cable develops aproperty of having a close relation between the degree of rotation atits proximal and distal ends. This allows the simplification of thedesign of the ultrasound catheter by making the angle of rotation at thedistal end of the torque cable (within the body) a reasonableapproximation of the angle of rotation at the proximal end of the torquecable (outside of the body).

The rotation of the torque cable or shaft at the point from which theimaging occurs may not be identical to the rotation occurs at theproximal end of the torque cable or shaft. This occurs especially whenthe flexible shaft is delivered through tortuous passageways and is, atleast in part, due to inertia and friction between the rotatingcomponents and stationary components of the imaging shaft. Theassumption that the rotational speed of the proximal and distal ends ofthe rotary shaft are equal to each other is also less likely to be validif the rotational speed varies over time. The undesirable result of notknowing the true angular velocity of the imaging probe at the point fromwhich the imaging beam is directed towards the tissue leads to anartifact referred to non-uniform rotational distortion (NURD). NURD canlead to significant distortion of the image and a concomitant reductionin the geometric accuracy of the image. Knowledge of a more preciseestimation of the true rotary speed of the distal rotary shaft or animaging assembly attached to the rotary shaft can help overcome suchdistortion by providing more accurate information for imagereconstruction. A better estimation of the rotary speed can also helpimprove the accuracy of co-registration of images when more than oneimaging modality is implemented on the imaging probe (such as combinedultrasound and optical imaging).

While use of more than one type of imaging technique, such ultrasoundand optical techniques, have both proved valuable in medicalapplications for high resolution imaging, they are not commonly used intandem. As described in the Summary of the related art below, there aresome designs that exist for the combination of optical and ultrasoundtechnologies. However, the limitations in these designs have preventedtheir acceptance.

Namely, designs that incorporate optical and ultrasound technologiesoffset the ultrasound and optical imaging mechanisms, such as disclosedin (Maschke, U.S. Pat. No. 7,289,842 resulting in the acquisition ofunaligned ultrasound and optical signals. Alignment of the resultantdata from these two imaging means requires movement of the imagingmechanisms and is prone to registration errors due to (i) non-uniformrotational distortion (NURD), (ii) motion of the object occurringbetween successive imaging of the same location using the two imagingmeans, (iii) variability in the object being imaged, and (iv) difficultyin accurately tracking the location of the imaging means. All theseeffects result in inaccurate co-registration which limits the usefulnessof the acquisition of data from the two imaging means.

SUMMARY OF THE RELATED ART

A catheter-based system for intravascular ultrasound is described byYock (U.S. Pat. No. 4,794,931) to provide high resolution imaging ofstructures in blood vessels. This system comprises an outer sheath,within which there is an ultrasound transducer near the distal end of along torque cable. When a motor rotates the torque cable and ultrasoundtransducer assembly, 2D cross-sectional images of anatomical structures,such as blood vessels, can be made. Linear translation of the catheteror the torque cable and ultrasound transducer in combination with therotational motion of the ultrasound transducer allows for acquisition ofa series of 2D images along the length of the catheter.

Milo et al (U.S. Pat. No. 5,429,136) and Lenker et al (U.S. Pat. Nos.6,110,121 and 6,592,526) describe reciprocating and vibrating means forscanning an ultrasound imaging beam in circumferential or longitudinaldirections at the end of the catheter. Reciprocating or vibrating meansobviates the need to use a mechanism such as a slip ring to provide anelectrical connection to a probe that rotates more than a few rotationsin a particular direction, such as more than one or two rotations.Similarly, certain implementations of optical imaging can avoid the useof optical rotary joints using reciprocating or vibrating means.

Liang et al. (U.S. Pat. Nos. 5,606,975 and 5,651,366) describe means ofimplementing forward-looking intravascular ultrasound where ultrasoundis directed towards a mirror that causes the ultrasound beam topropagate at an angle from the longitudinal axis of a rotating torquecable advanced within the vasculature. Liang et al. also describe meansof varying the angle of deflection of the mirror using either amicromotor, a gear clutch mechanism, steering cables or bimorph elementssuch a shape memory alloys, piezoelectric files or conductive polymers.FIG. 13 of U.S. Pat. No. 5,651,366 shows a diagram of a forward lookingultrasound probe combined with a fiber optic to deliver laser ablationenergy via a fiber and mirror in a coaxial direction to the ultrasoundimaging beam, but does not relate to combined optical and acousticimaging or provide for optical focusing elements which would be ofbenefit for imaging purposes.

The use of intravascular ultrasound (IVUS) has since become commonplace,with many improvements and adaptations to the technology. A flexibletorque cable (Crowley, U.S. Pat. No. 4,951,677) improves the fidelity ofthe transmission of rotational torque along the length of an IVUScatheter, minimizing an artifact known as non-uniform rotationaldistortion.

The center frequency of IVUS lies within the range of 3 to 100 MHz andmore typically in the range of 20 to 50 MHz. Higher frequencies providehigher resolution but result in worse signal penetration and thus asmaller field of view. Depth of penetration can range from less than amillimeter to several centimeters depending on several parameters suchas center frequency and geometry of the transducer, the attenuation ofthe media through which the imaging occurs and implementation-specificspecifications that affect the signal to noise ratio of the system.

Variations of high frequency ultrasound exist, where the signalacquisition and/or analysis of the backscattered signal is modified tofacilitate obtaining or inferring further information about the imagedtissue exist. These include elastography, where the strain within tissueis assessed as the tissue is compressed at different blood pressures (deKorte et al Circulation. 2002 Apr. 9; 105(14):1627-30); Doppler imagingwhich assesses motion such as blood flow within anatomic structures;virtual histology, which attempts to infer the composition of tissueusing the radio-frequency properties of the backscattered signalcombined with a pattern recognition algorithm (Nair, U.S. Pat. No.6,200,268); second harmonic imaging (Goertz et al, Invest Radiol. 2006August; 41(8):631-8) and others. Each of these forms of imaging can beimproved upon by means described in the present invention.

Ultrasound transducers themselves are improving considerably, includingthe use of single crystal ultrasound transducers and compositeultrasound transducers.

Hossack et al (WO/2006/121851) describe a forward looking ultrasoundtransducer using a CMUT transducer and a reflective surface.

Tearney et al (U.S. Pat. No. 6,134,003) describe several embodimentsthat enable optical coherence tomography to provide higher resolutionimaging than is readily obtained by high frequency ultrasound or IVUS.

Boppart et al (U.S. Pat. No. 6,485,413) describe several embodiments ofoptical coherence tomography imaging, including forward-lookingimplementations. Either an optical fiber or a gradient index (GRIN) lensare displaced using a mechanism such as a motor, a piezoelectric, amoveable wire, inflation means and others.

Mao et al (Appl Opt. 2007 Aug. 10; 46(23):5887-94) describe methods forcreating ultrasmall OCT probes using single mode fiber, coupled to asmall length of GRIN fiber which acts as a lens. Including an opticalspacer between the fiber and the lens can alter the working distance ofthe fiber-lens system. Furthermore, adding a small length of no-cladfiber to the distal end, and cutting the no-clad fiber at an angle canadd a deflecting element to the end of the fiber-lens system. Thisdeflecting element enables side-viewing imaging, which could also beaccomplished using a small prism or mirror.

Variations of optical coherence tomography (OCT) include polarizationsensitive OCT (PS-OCT) where the birefringent properties of tissuecomponents can be exploited to obtain additional information aboutstructure and composition; spectroscopic OCT which similarly providesimproved information regarding the composition of the imaged structures;Doppler OCT which provides information regarding flow and motion;elastography via OCT; and optical frequency domain imaging (OFDI), whichallows for a markedly more rapid acquisition of imaging data andtherefore enables imaging to occur over a larger volume of interest inless time. Again, each of these forms of imaging can be improved upon bymeans of the present invention.

Several other forms of fiber-optic based imaging exist other than OCT.Amundson et al describe a system for imaging through blood usinginfra-red light (U.S. Pat. No. 6,178,346). The range of theelectromagnetic spectrum that is used for their imaging system isselected to be one which optimizes penetration through blood, allowingoptical imaging through blood similar to that afforded by angioscopy inthe visible spectrum, but without the need to flush blood away from theregion being imaged.

Dewhurst (U.S. Pat. No. 5,718,231) discloses a forward looking probe forintravascular imaging where a fiber optic travels through an ultrasoundtransducer to shine light on a target tissue straight in front of theend of the probe. The light then interacts with the target tissue andmakes ultrasound waves, which are received by the ultrasound sensor andthe images are photoacoustic images only as the system is not configuredto receive and process optical images. The ultrasound sensor used in theDewhurst device is limited to thin film polymeric piezoelectrics, suchas thin film PVDF, and is used only to receive ultrasound energy, not toconvert electrical energy to ultrasound.

Angioscopy, endoscopy, bronchoscopy and many other imaging devices havebeen described which allow for the visualization of internal conduitsand structures (such as vessels, gastrointestinal lumens and thepulmonary system) in mammalian bodies based on the principle ofilluminating a region within the body near the distal end of a rigid orflexible shaft. Images are then created by either having a photodetectorarray (such as a CCD array) near the end of the shaft or by having abundle of fiber optics transmit the received light from the distal endof the shaft to the proximal end where a photodetector array or othersystem that allows the operator to generate or look at an imagerepresentative of the illuminated region. Fiber bundles are bulky andreduce the flexibility of the shaft among other disadvantages.

Other fiber optic based modalities for minimally invasive assessment ofanatomic structures include Raman spectroscopy as described by Motz etal (J Biomed Opt. 2006 March-April; 11(2)), near infrared spectroscopyas described by Caplan et al (J Am Coll Cardiol. 2006 Apr. 18; 47(8Suppl):C92-6) and fluorescence imaging, such as tagged fluorescentimaging of proteolytic enzymes in tumors (Radiology. 2004 June;231(3):659-66).

The ability to combine ultrasound and optical coherence tomography ontoa single catheter would be extremely advantageous. Kubo et al presentedan interesting in vivo study of coronary arteries using OCT, IVUS andangioscopy to assess the morphology of lesions that have caused an acutemyocardial infarction (Journal of American College of Cardiology, Sep.4, 2007, 10(50):933-39). They demonstrate that there are benefits toimaging with each of these modalities. However, in order to executetheir study, they had to use separate catheters for each of IVUS, OCTand angioscopy imaging modalities as no catheters that combine thesefunctions have been commercialized to date. Kawasaki et al previouslycompared OCT, conventional IVUS and a variant of IVUS known asintegrated backscatter IVUS on cadaveric specimens of coronary arteriesusing separate probes for the OCT and IVUS components. Brezinski et al(Heart. 1997 May; 77(5):397-403) had previously demonstrated ex vivostudies on dissected aortic specimens where IVUS and OCT images werecompared, again using separate probes. The OCT probes in this latterstudy were not suitable for in vivo use.

Optical coherence tomography generally has superior resolution toultrasound and has the potential to better identify some structures orcomponents in vascular and other tissues than ultrasound. For example,fibrous cap thickness or the presence of inflammatory or necroticregions near the surface of arteries may be better resolved with opticalcoherence tomography. However, optical coherence tomography is limitedby its small penetration depth (on the order of 500 to 3000 microns) inmost biologic media. Most such media are not optically transparent.

Meanwhile, ultrasound has the ability to better penetrate throughbiological media such as blood and soft tissues and has a depth ofpenetration that typically extends several millimeters or centimetersbeyond that of optical coherence tomography. The ability to image witheither or both methods of imaging using a combined imaging deviceprovides advantages with respect to selecting the required resolutionand depth of penetration. Furthermore, much of the information acquiredby optical coherence tomography is complementary to that acquired byultrasound and analysis or display of information acquired by bothimaging methods would improve the ability to better understand theinterrogated tissue, such as with respect to its composition.

These differences between IVUS and OCT are well known in the art.Maschke (United States Patent Publication No. 2006/0116571 correspondingto U.S. patent application Ser. No. 11/291,593) describes an embodimentof a guidewire with both OCT and IVUS imaging transducers mounted uponit.

The described invention has several shortcomings. Guidewires aretypically 0.014″ to 0.035″ in diameter (approximately 350 microns to 875microns), yet ultrasound transducers typically are at least 400microns×400 microns and generally are larger in size for the frequenciesin the 20 to 100 MHz range. If the transducer is too small, the beam ispoorly focused and has poor signal properties. In Maschke the IVUS andOCT imaging mechanisms are located at different positions along thelength of the guidewire and a drawback to this type of configurationhaving the IVUS and OCT imaging means located at different positionsalong the length of an imaging shaft does not allow for optimalco-registration of images.

U.S. Pat. No. 7,289,842) issued to Maschke describes an imaging systemthat combines IVUS and OCT on a catheter where the IVUS and OCT imagingelements are longitudinally displaced from each other along the lengthof a catheter that rotates around its longitudinal axis. Maschke alsodescribes generating images where the center portion of the images aresubstantially derived from the output of the higher resolution OCTimaging portion of the system while the outer portion of the images aresubstantially derived from the output of the ultrasound imaging portionof the system, to make use of ultrasound's greater depth of penetrationin combination with OCT's higher resolution for tissues close to thecatheter.

Park et al (U.S. patent application Ser. No. 11/415,848) also brieflyrefers to the notion of having a catheter that combines IVUS and OCTimaging onto a single catheter.

However, the integration of means for combined acoustic and opticalimaging, such as combined IVUS and OCT imaging, onto a single device isnot trivial. Having an optical imaging element and an acoustic imagingelement longitudinally separated from each other on a primarily rotatingcatheter does not provide an ideal configuration for combined imaging. Amore ideal configuration would enable the acquisition of high qualityacoustic and optical signals from which ultrasound and optical-basedimages could be made while enabling the acoustic and optical images tobe registered with each other in a highly precise manner.

For example, by simply placing an IVUS imaging element in line with anOCT imaging element along the length of the catheter, the center of theimaging planes of the IVUS and OCT images will be separated from oneanother by a distance of at least approximately half the length of theultrasound transducer and half the length of the optical imagingelements.

Mechanical IVUS transducers for vascular imaging are typically more than400 microns in length. The separation between the IVUS and OCT planes ofimaging in a configuration such as that proposed by Maschke wouldrequire at least 250 microns of separation between the optical andacoustic imaging planes. Typically, mechanical IVUS rotates at 30 framesper second with a pullback rate of 0.5 mm/s, meaning that from a giventime point to, at least 15 imaging frames or 500 milliseconds wouldelapse between the time that the more distally placed imaging meanswould translate to the same position at which the more proximally placedimaging means was originally positioned at time to. This separation ofseveral hundred milliseconds or several rotations of the imaging probemakes it difficult to precisely register the imaging data from oneimaging means with the other.

This is particularly relevant given the fact that the catheter canundergo significant unintentional lateral and longitudinal displacementswithin body lumen in that time period, such as those displacements thatoccur as a result of cardiac contraction and pulsatile flow. Non-uniformrotational distortion (NURD) can also have an impact on the ability toaccurately register images acquired several rotations apart from eachother. Any imprecision of the registration of the two data sets is evenmore significant when one considers the scale at which importantpathologies, such as vulnerable plaques can be found. Dramaticdifferences in the appearance of an arterial plaque's composition (e.g.the thickness of a fibrous cap, the presence of a calcified nodule orthe extent of an atheromatous deposit) can be observed in as little as afew hundreds microns along the length of a vessel. Similarly, small butpotentially relevant sidebranches of anatomic conduits, such as bloodvessels, can have dimensions on the order of less than a hundredmicrons.

Previous experiments and implementations of IVUS and OCT or othercombinations of acoustic and optical imaging have not been provided thatenable significant precision in the registration of the imaging datafrom the two or more imaging means in a manner that is suitable forminimally invasive imaging, such as intravascular imaging.

To the best of our knowledge, previous experiments and implementationsof IVUS and OCT or other combinations of acoustic and optical imaginghave not been provided that enable significant precision in theregistration of the imaging data from the two or more imaging means in amanner that is suitable for minimally invasive imaging, such asintravascular imaging.

It would be very advantageous to also provide high resolution imagingprobes that combine acoustic and optical imaging onto “forward-looking”probes rather than “side-viewing” probes. It would also be helpful toprovide similar probes that can look backwards, or from multiple anglesin a generally side-viewing configuration.

It would also be advantageous to provide high-resolution imaging probesthat combine ultrasound imaging with one or more optical imaging means.

It would also be advantageous to provide minimally invasive imagingprobes that can be used for photoacoustic imaging or sonoluminescentimaging.

It would also be advantageous to provide minimally invasive imagingmeans where on of the imaging means provides helpful informationregarding the direction in which the other imaging means is acquiringimaging data.

SUMMARY OF THE INVENTION

The present invention provides embodiments of imaging probes forcombining acoustic and optical imaging means in a manner thatfacilitates simultaneous imaging by two or more imaging methods. Theembodiments enable methods to accurately co-register the images obtainedfrom each of the modalities. In some embodiments, the current inventionprovides embodiments for combining acoustic imaging means with thedelivery of therapeutic energy, such as ultraviolet light forphotodynamic therapy or laser energy for ablation procedures.

The present invention also provides embodiments where one form ofimaging is used to help with the reconstruction of the second form ofimaging. This is more specifically related to monitoring the position ororientation of a component in the image probe that subsequentlydetermines the position or orientation of the imaged region.

The present invention provides methods for combining high frequencyultrasound and optical coherence tomography into a combined imagingsystem.

The present invention provides novel means for implementing a combinedultrasound and optical imaging system where the volume scanned includesa region either forward of, or behind, the location of the imagingtransducers.

The present invention provides the ability to take images similar tothose produced by angioscopy, endoscopy and similar imaging techniquesusing a single optic or a small number of fiber optics, in combinationwith means to acquire ultrasound images. These optical images can alsobe acquired using infrared and/or visible wavelengths.

The present invention provides means for combining high frequencyultrasound and optical coherence tomography where the volumes scannedinclude regions either forward of, or behind, the locations of theimaging transducers.

Embodiments of the present invention are able to scan a region for thepurposes of imaging or delivery of therapeutic energy the regionaccessed by a shaft where changes in the rotation velocity of the shaftcauses changes in the direction of either an emitter and/or receiver ofacoustic and/or optical energy.

The present invention also facilitates certain forms of high resolutionimaging that use acoustic energy to create optical energy(sonoluminescence imaging) or optical energy to create acoustic energy(photoacoustic imaging).

An embodiment of the present invention provides an imaging probe forinsertion into bodily lumens and cavities for imaging an interior ofsaid bodily lumens and cavities or imaging exterior surfaces of a body,comprising:

a) an hollow shaft having a longitudinal axis having distal and proximalend sections and an midsection, an imaging assembly being located insaid hollow shaft, said imaging assembly being connected to a first endof an imaging conduit, said imaging conduit extending through the hollowshaft and being connectable at a second end thereof to an imageprocessing and display system through the proximal end section, saidimaging conduit including a fiber optic having a distal end and saidimaging assembly including an optical emitter/collector including lightdirecting and receiving means associated with said distal end of a fiberoptic for directing light imaging energy out of a distal end of saidfiber optic and receiving reflected light imaging energy signals anddirecting said received reflected light energy signals back to saidimage processing and display system, said imaging assembly including anultrasound transducer and said ultrasound transducer emitting andreceiving reflected ultrasound imaging energy signals and said imagingconduit including an electrical conductor for electrically coupling theultrasound transducer to an ultrasound signal generator connectable tosaid second end of said imaging conduit, said imaging conduit beingconnectable at said second end to a source of light;

b) said imaging assembly including a scanning mechanism configured todeliver said light from the optical emitter/collector and ultrasoundfrom said ultrasound transducer along pre-selected paths out of saidhollow shaft, the ultrasound transducer and the opticalemitter/collector being positioned and oriented relative to each otherto enable accurate co-registering of received reflected light imagingenergy signals and reflected ultrasound imaging energy signals duringscanning a region of interest; and

c) drive mechanism for imparting motion to said imaging conduit and saidimaging assembly, said drive mechanism being connectable to a controllerwhich is connectable to said image processing and display system.

A further understanding of the functional and advantageous aspects ofthe invention can be realized by reference to the following detaileddescription and drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

Preferred embodiments of the invention will now be described, by way ofexample only, with reference to the drawings, in which:

FIG. 1 is a schematic of an imaging system including ultrasound andoptical imaging components;

FIG. 2 is a perspective drawing of a flexible imaging probe with anadapter, conduit and imaging assembly;

FIG. 2 a is a cross sectional view of the mid section of the imagingprobe of FIG. 2 taken along the dotted line;

FIG. 2 b is an expanded perspective drawing of the distal region of theimaging probe of FIG. 2;

FIG. 2 c shows a schematic of how the rotary and non-rotary componentsof the imaging probe can be coupled with an adapter to the rest of animaging system.

FIG. 2 d is a perspective drawing of an example of the coupling of therotary and non-rotary components of the probe to an adapter.

FIGS. 3 a to 3 e are representative of general imaging catheterconfigurations described in the prior art;

FIG. 3 a shows one embodiment of an over-the-wire configuration for anexternal sheath that may be incorporated with the imaging probe if aguidewire lumen is included;

FIG. 3 b shows a cross-section through the imaging probe to demonstratethe guidewire lumen configuration.

FIG. 3 c shows a rapid access configuration for an external sheath thatmay be incorporated with the imaging probe if a guidewire lumen isincluded;

FIG. 3 d shows a cross-section through a portion of the imaging probethat does not contain a guidewire lumen;

FIG. 3 e shows a cross-section through a portion of the imaging probethat does contain a guidewire lumen;

FIGS. 4 a to 4 l are examples of ultrasound transducers that contain ahole for allowing transmission of optical energy through the transducerthat enables optical and acoustic imaging of regions that are preciselyaligned with each other, as well as means to deflect the path of theimaging light;

FIGS. 5 a to 5 f are examples of ultrasound transducers that contain ahole for allowing transmission of optical energy through the transducerthat enables optical and acoustic imaging of regions that are preciselyaligned with each other, without a means to deflect the path of theimaging light;

FIGS. 6 a to 6 c demonstrate representative acoustic transducerconfigurations, with FIG. 6 a not having a hole in the transducer. FIGS.6 d to 6 f demonstrate representative simulation results of the effectsof placing a hole through an ultrasound transducer on the acoustic beampattern produced by the ultrasound transducer with FIG. 6 d not having ahole;

FIGS. 7 a to 7 e show examples of ultrasound transducers that have anoptical apparatus for transmitting and/or receiving optical imagingenergy either on top of or recessed within an acoustic transducer;

FIG. 8 a is a perspective view of an imaging assembly suitable for sideviewing with both acoustic and optical imaging;

FIG. 8 b is a side view of the imaging assembly in FIG. 8 a;

FIGS. 8 c to 8 e are end views of the imaging assembly in FIG. 8 a indifferent rotated positions;

FIGS. 9 a to 9 c depict configurations whereby an optical imagingemitter/receiver is embedded into the backing material 435 of anacoustic transducer.

FIGS. 10 a to 10 e are similar to FIGS. 8 b to 8 e showing the imagingassembly being rotated in a reciprocating fashion rather than in asingle rotational direction;

FIG. 11 shows a perspective view of an imaging probe where thepredominant motion is a longitudinal motion where the surface swept bythe optical beam and the acoustic beam are two co-planar rectangles;

FIG. 12 shows a perspective view of an embodiment of an imaging probewhere the optical imaging system is configured such that the opticalimaging beams are angled such that these imaging beams substantiallyconverge or overlap;

FIG. 13 is a cross sectional view of an imaging assembly suitable forside viewing with both acoustic and optical imaging;

FIG. 14 a is a cross sectional view of an imaging assembly suitable forforward viewing with both acoustic and optical imaging;

FIG. 14 b is a cross sectional view of an imaging assembly suitable forforward viewing with both acoustic and optical imaging in which anartificial muscle polymer can be used to deform the distal region of theimaging probe;

FIG. 15 a is a cross sectional view of an imaging assembly suitable forside viewing with both acoustic and optical imaging using a reflectivecomponent to direct the optical and acoustic beams in the sidewaysdirection;

FIGS. 15 b and 15 c are similar to FIG. 15 a but in which the reflectivecomponent is mounted about a pivot point so the optical and acousticbeams can be scanned in the sideways direction at a variable angle;

FIG. 16 a is a cross section of an embodiment of an imaging probe usinga tiltable component where the tilting action is modulated bycentripetal acceleration due to the rotational motion of the imagingassembly around the longitudinal axis;

FIG. 16 b is a view along the line 16 b-16 b of FIG. 16 a;

FIG. 16 c is a cross section of the imaging probe of FIG. 16 a but withthe tiltable component at a different angle during use;

FIG. 16 d is a view along the line 16 d-16 d of FIG. 16 c;

FIG. 17 a is a perspective drawing of a deflecting component thatcomprises a flat optically reflective layer and a shaped acousticallyreflective layer;

FIGS. 17 b through 17 d depict cross-sections of the deflectingcomponent of FIG. 17 a;

FIG. 18 a is a perspective view of an ultrasound imaging transducer withtwo (2) optical imaging emitters/receivers through two (2) separateoptically transmissive channels in the acoustic transducer;

FIG. 18 b is a perspective view of an embodiment of an imaging probehaving an ultrasound imaging transducer with two (2) optical imagingemitters/receivers arranged in a manner such that they are aligned withthe predominant rotary motion of the imaging assembly;

FIG. 18 c is a view along arrow C of FIG. 18 b;

FIG. 19 is a schematic of a system where there are two optical imagingsystems that are coupled to the same optical imaging waveguide viaoptical routing circuitry;

FIGS. 20 a and 20 b demonstrate sector-shaped patterns forsimultaneously demonstrating portions of two (2) or more images that areco-registered with each other;

FIGS. 21 a and 21 b demonstrate arbitrary patterns for simultaneouslyshowing portions of 2 or more images that are co-registered with eachother;

FIG. 22 is a schematic of a display which transitions over time from oneimage to another, co-registered image;

FIGS. 23 a and 23 b demonstrate how a feature in a first image can bemapped onto a feature in another image that is co-registered with thefirst image;

FIGS. 24 a and 24 b demonstrate how a contour feature in a first imagecan be mapped into another image with is co-registered with the firstimage and vice versa;

FIGS. 25 a and 25 b provide a schematic for how a composite image can beconstructed from two (2) or more co-registered imaging datasets; and

FIGS. 26 a to 26 c shows cross sectional views of an imaging probe witha rotary encoder.

DETAILED DESCRIPTION OF THE INVENTION

Without limitation, the majority of the systems described herein aredirected to an imaging probe that enables imaging by both optical andacoustic means. As required, embodiments of the present invention aredisclosed herein. However, the disclosed embodiments are merelyexemplary, and it should be understood that the invention may beembodied in many various and alternative forms.

The Figures are not to scale and some features may be exaggerated orminimized to show details of particular elements while related elementsmay have been eliminated to prevent obscuring novel aspects. Therefore,specific structural and functional details disclosed herein are not tobe interpreted as limiting but merely as a basis for the claims and as arepresentative basis for teaching one skilled in the art to variouslyemploy the present invention. For purposes of teaching and notlimitation, the illustrated embodiments are directed to an imaging probethat enables imaging by both optical and acoustic means.

As used herein, the term “about”, when used in conjunction with rangesof dimensions, temperatures or other physical properties orcharacteristics is meant to cover slight variations that may exist inthe upper and lower limits of the ranges of dimensions so as to notexclude embodiments where on average most of the dimensions aresatisfied but where statistically dimensions may exist outside thisregion. For example, in embodiments of the present invention dimensionsof components of an imaging probe are given but it will be understoodthat these are not meant to be limiting.

As used herein, the phrase “co-registration of images” refers to theprocess of identifying a subset of imaging data acquired by one imagingmeans with a subset of imaging data acquired using another imaging meanswhere the identified imaging data from the two means was acquired bydetecting a form of imaging energy (e.g. photons or ultrasound) from thesame object (or tissue in the case of the present invention). Eachco-registered point in the first subset can then be mapped to acorresponding point in the second subset such that the two points fromthe two different imaging means are thought to have been acquired from asimilar focal region of the imaged object (or tissue).

Successful and accurate co-registration of images, or portions thereof,between images acquired using two (2) or more imaging means is helpfulin that it can provide multiple opportunities to assess features ofinterest of the imaged object by more than one imaging means.

FIG. 1 represents an overview of an exemplary imaging system constructedin accordance with the present invention shown generally at 10. Itcomprises an imaging probe 12, which connects via an adapter 14 to animage processing and display system 16. The image processing and displaysystem 16 comprises the necessary hardware to support one or more of thefollowing imaging modalities: 1) ultrasound, 2) optical coherencetomography, 3) angioscopy, 4) infrared imaging, 5) near infraredimaging, 6) Raman spectroscopy-based imaging and 7) fluorescenceimaging.

Implementations of the optical coherence tomography, ultrasound,angioscopy and infrared imaging circuitry have been described in theprior art.

The system herein described further typically comprises a controller andprocessing unit 18 to facilitate the coordinated activity of the manyfunctional units of the system, and may further comprise a displayand/or user interface and may further comprise electrode sensors toacquire electrocardiogram signals from the body of the patient beingimaged. The electrocardiogram signals may be used to time theacquisition of imaging data in situations where cardiac motion may havean impact on image quality. The optical circuits and electronics 21forming image processing and display system, if included in a particularimplementation of the present invention, may include any or all of thefollowing components: interferometer components, one or more opticalreference arms, optical multiplexors, optical demultiplexors, lightsources, photodetectors, spectrometers, polarization filters,polarization controllers, timing circuitry, analog to digital convertersand other components known to facilitate any of the optical imagingtechniques described in the background and prior art sections. Theultrasound circuitry 20 may include any or all of the followingcomponents: pulse generators, electronic filters, analog to digitalconverters, parallel processing arrays, envelope detection, amplifiersincluding time gain compensation amplifiers and other components knownto facilitate any of the acoustic imaging techniques described in thebackground and prior art sections.

The controller and processing units 18, if included in a particularimplementation of the present invention, serve multiple purposes and thecomponents would be markedly adapted based on the needs of a particularimaging system. It could include one or a combination of motor drivecontroller, data storage components (such as memory, hard drives,removable storage devices, readers and recorders for portable storagemedia such as CDs and DVDs), position sensing circuitry, timingcircuitry, cardiac gating functionality, volumetric imaging processors,scan converters and others. A display and user interface 22 is alsooptionally provided for either real time display or display of data at atime later than the time at which imaging data is acquired.

The imaging probe 12 comprises an imaging assembly 30 near its distalend 32, an optional conduit 34 along a substantial portion of itslength, and a connector 36 at its proximal end 38. For the purposes ofthis invention, an imaging assembly 30 generally refers to the componentof the imaging probe 12 from which the signals (acoustic or optical (orboth)) are collected for the purposes of imaging a region that isproximate to the imaging assembly 30. The imaging assembly 30 includesat least one or more emitters of imaging energy and at least one or morereceivers of imaging energy. For the purposes of this invention,“imaging energy” refers to both light and acoustic energy. Specifically,light refers to electromagnetic waves that span the ultraviolet, visibleand infrared spectrum of wavelengths. For example, for acoustic imaging,the imaging assembly 30 contains an ultrasound transducer that is bothan emitter and receiver of acoustic energy.

For optical imaging, the imaging assembly 30 typically contains thedistal tip of a fiber optic, as well as a combination of opticalcomponents such as a lens (such as a ball lens or GRIN lens), whichcollectively serve the purpose of acting as an optical receiver and mayalso serve as an optical emitter. A mirror and/or a prism are oftenincorporated as part of an optical emitter and/or receiver. The imagingassembly 30, connector 36 and/or imaging conduit 34 may beliquid-filled, such as with saline and may be flushed.

The imaging probe 12 may contain ports at one or more points along itslength to facilitate flushing. For optical imaging, it is possible toconsider a gas filled imaging probe 12. Preferably, the gas wouldsubstantially comprise carbon dioxide or another readily dissolved gas.Alternatively, the imaging assembly may be compartmentalized such thatthere is at least one gas-filled compartment or lumen for opticalimaging and at least one fluid-filled compartment or chamber foracoustic imaging.

The imaging conduit 34 comprises at least one optical waveguide and atleast one conductive wire (preferably two or more) that connect anemitter and/or receiver via a connector to an adapter. The imagingconduit 34 may also act as a mechanical force transmission mechanism forrotating or translating the imaging assembly. For example, the imagingconduit 34 may comprise a fiber optic, wrapped by two layers ofelectrical wire that are insulated by each other. The imaging conduit 34may further be reinforced by other structural features, such ashelically wrapped wires or other designs used to construct imagingtorque cables for rotating scan mechanisms, as described in the priorart.

The adapter 14 facilitates transmission of signals within any fibersand/or wires to the appropriate image processing units. The adapter 14may also incorporate a pullback mechanism 49 (FIG. 2 d) or areciprocating push-pull mechanism to facilitate longitudinal translationof the imaging assembly. Such longitudinal translation of the imagingassembly 30 may occur in conjunction with the longitudinal translationof an external shaft that surrounds the imaging conduit 34, or may occurwithin a relatively stationary external shaft.

Additional sensors may be incorporated as part of the adapter 14, suchas position sensing circuitry, for example to sense the angle ofrotation of a rotary component within the imaging probe 12. The imagingprobe 12 may also include a memory component such as an EEPROM or otherprogrammable memory device that includes information regarding theimaging probe to the rest of the imaging system. For example, it mayinclude specifications regarding the identification of specifications ofthe imaging probe 12 and may also include calibration informationregarding the probe 12.

While precise alignment of the acoustic and optical imaging data ishighly desired, it is also important to recognize the need to optimizethe geometry of a minimally invasive probe so that it is as small asreasonably possible to achieve its desired purpose. Current IVUS probesare approximately 0.9 to 2 mm in diameter and the smaller sizes ofprobes can be delivered more distally within the vascular tree of thecoronary anatomy as the vessel size tapers down. Thus, smaller sizesgenerally allow for interrogation of a larger portion of the coronaryanatomy. It is therefore desirable to have embodiments of a probe thatcombines optical and acoustic imaging in arrangements that minimizecertain dimensions of the probe, such as the diameter of the probe.

FIG. 2 is a perspective drawing of a flexible catheter containing afiber optic 40 and a co-axial electrical wire 50. The proximal connectorcontains fiber optic 40 that can be received by the adapter to opticallycouple the imaging fiber optic 40 to the optical imaging system“back-end”. There are also electrical connectors 56 that allow the oneor more electrical conduits to be connected to the ultrasound circuitry20 and/or controller and processing units 18. In embodiments where theimaging conduit rotates around its longitudinal axis, there may be aneed to couple the rotating components of the imaging fiber optic withthe relatively stationary fiber optic that connects to the opticalimaging system's back-end 21. The coupling of a rotating fiber opticprobe can be accomplished using a fiber optic rotary joint incorporatedeither as part of the proximal connector of the imaging probe 10 or aspart of the adapter 14. Similarly, in embodiments where the imagingconduit rotates around its longitudinal axis, there may be a need tocouple the conductive wires that rotate with the imaging conduit withthe relatively stationary conductors of the ultrasound circuitry 20and/or controller and processing units 18, preferably by means of sliprings. These slip rings can be incorporated as part of the proximalconnector of the imaging probe 36 or as part of the adapter 14.

FIG. 2 a shows a cross sectional view of the mid section of the imagingprobe of FIG. 2 taken along the dotted line which shows a fiber optic40, guidewire port 44 and guide wire 42, imaging conduit 34, imagingconduit lumen 46, external sheath 48 which is a hollow, flexibleelongate shaft made of a physiologically compatible material and havinga diameter suitable to permit insertion of the hollow elongate shaftinto bodily lumens and cavities, and coaxial electrical wiring 50. Theexpanded detailed view of the end of the imaging probe 10 shown in FIG.2 b shows the distal end of the guidewire 42 extended beyond the end ofthe outer sheath 48 and a flush port 54 at the end of the sheath 48. InFIG. 2 the proximal end of the imaging probe 10 includes anotherguidewire port 55 into which guidewire 42 is inserted and the connectorassembly 36 which includes a flush port 58 and electrical contacts 56along the connector body.

FIG. 2 c shows a schematic of how the rotary and non-rotary componentsof the imaging probe can be coupled with an adapter to the rest of animaging system. FIG. 2 d schematically shows how the rotating componentsof the imaging probe can be coupled to the rotating components of anadapter. The rotating components of each can be electrically, opticallyand/or mechanically coupled using connectors and other configurationsknown in the art. Similarly, the non-rotating components of the imagingprobe can be coupled to the non-rotating components of the adapter 14.The adapter 14 can include slip rings, optical rotary joints and othersuch implements for electrically or optically coupling a rotarycomponent to a non-rotary component and enable communication ofnecessary electrical and optical signals with the rest of the system.

Dual-fiber optical rotary joints are also available but considerablymore complex. Electrical coupling between any conductor mounted onto arotating component in the imaging probe 12 can be coupled tonon-rotating conducting elements via metallic slip rings and springs,metallic slip rings and brushes or other commonly known methods offorming conductive contact between a stationary conductor and a rotaryconductor.

While the electrical, optical and mechanical connections are shownseparately in FIG. 2 d, it is possible to reduce the several connectorsthat must each be separately connected between the probe and adapterwith fewer connectors by combining several connectors into combinedconnectors, as needed for a specific embodiment.

FIG. 3 a shows one embodiment of an over-the-wire configuration for anexternal sheath at 47 and FIG. 3 b shows a cross-section of sheath 47through the portion that contains the imaging assembly 30 along thevertical line 3 b-3 b in FIG. 3 a.

FIG. 3 c shows an embodiment at 60 that is a “rapid exchange”configuration for the external sheath that may be incorporated with theimaging probe if a guidewire is required. Sheath 60 in FIG. 3 c includesthe entry port 55 shown in FIG. 2. FIG. 3 d shows a cross-section of the“rapid-exchange” configuration 60 through the portion that is proximalto the entry port 55 for a guidewire along line 3 d-3 d in FIG. 3 c.FIG. 3 e shows a cross-section along line 3 e-3 e in FIG. 3 c.

The present invention describes several embodiments by which preciselyregistered ultrasound and optical images can be formed. The simplestconceptual approach is to have the paths of the ultrasound and opticalimaging beams be aligned collinearly with each other.

Referring to FIG. 4 a, an imaging probe 399 is provided which isconfigured to allow imaging by acoustic and optical means in the samedirection, so that an acoustic transducer that allows light energy totravel through a channel in the transducer is utilized. Essentially,probe 399 uses an acoustic transducer 402 that is altered to have anoptically transmissive channel made through its substrate. The acoustictransducer 402 can be any kind of ultrasound transducer known in theart, such as piezoelectric composition (e.g. PZT or PVDF), a compositetransducer or a single crystal transducer.

Electrical conductors 400 are directed to the conducting layers 401 oneither side of the transducer's acoustic substrate 402. A fiber optic403 provides an optical conduit for enabling optical imaging. One ormore matching layers can be added to the emission surfaces of thetransducer, such as an epoxy layer (such as a silver or copperconductive epoxy layer which may functionally also serve as one or bothof the electrodes that drives the transducer), or a polymer (such asparylene or PVDF).

The optically transmissive channel 407 is made by any of severaltechniques, such as precision drilling, laser ablation, photo-etching,inclusion of a feature in a mold to create the opening and others.Precision drilling may include the use of drill bits, such as diamond orcarbide drill bits explicitly designed for cutting through hardmaterials. A high precision spindle, such as an air spindle, may behelpful for accurate and efficient execution of the drilling technique.A laser source can be used to ablate a channel through the substrate.Exemplary laser sources include YAG or excimer lasers.

Alternatively, if the acoustic transducer 402 is formed from a substratethat is initially viscous, a sacrificial component can be embedded inthe piezoelectric during the formation of the piezoelectric transducer402. The sacrificial component can then be removed by mechanical meansor exposure to a solvent. For example, a polystyrene cylinder can serveas the sacrificial component, which can be subsequently sacrificed usingdissolution in acetone. Alternatively, if the piezoelectric material 402is formed from a substrate that is initially viscous, a removablemandrel can be included in the material during the formation of thepiezoelectric transducer and removed after the piezoelectric haspartially or substantially hardened.

Conductive layers 401 on either side of the piezoelectric material 402are incorporated as required for applying a voltage to thepiezoelectric. The opening 407 is coupled to an optical waveguide 403,either directly, or by means of one or more mirrors 404 or prisms 397and one or more lenses 405. If any optical components are includedwithin the opening, a dampening, insulating layer of a compliantmaterial 406 (see FIG. 4 l), such as silicon or polymer may separate theoptical components from the acoustic substrate 402 to act as either anelectrical insulator or to minimize the transmission of stresses thatare generated by the acoustic substrate 402 to the optical components.

As in FIG. 4 b, the light from the fiber can be directed towards amirror 404 (or prism) that causes the light from the fiber to bedeflected through the optically transmissive channel 407. Alternatively,as in FIG. 4 c, a prism 397 can be used to deflect the light through theoptically transmissive channel. The prism 397 may deflect light eitheras a result of total internal reflection or be assisted by a reflectivecoating on its deflecting surface 419. The prism 397 may be a separateoptical component that is affixed to the appropriate position along theoptical path. For example, it can be glued in place onto the end of afiber, onto a lens or onto a spacer using bonding methods such as UVcured glue. Alternatively, attaching a no-clad optical fiber along theoptical path and cutting the segment of no-clad fiber at a desiredlength can be performed to make the prism. The segment of clad fiber canbe cut and/or polished to achieve the desired angle. Mao describes thismethod in the previously cited reference.

Also seen in FIG. 4 c, an optically transparent window 409 mayoptionally be found at the end of the optically transmissive channel 407and any unoccupied space within the channel may be filled with a gas,fluid or optically transparent material such as glass or any of severaltransparent polymers known in the art. The purpose of the window 409 isto prevent undesired air bubbles from being created or retained in thechannel 407 and to protect the components in the optically transmissivechannel 407.

As seen in FIG. 4 d it may be desirable to have a gas instead of fluidor solid material inside the channel 407 to improve the refractive powerof certain optical components such as a contoured lens 424, which may bea ball lens.

As seen in FIGS. 4 e to 4 g, the GRIN lens 405 or other opticalcomponent can reside adjacent to the distal dip of the optical fiber403, between the fiber 403 and the deflecting mirror or prism 397 alongthe optical path. In this case, the opening 407 in the acousticsubstrate 402 can be left free of any optical components and simplycontain an optically transparent material, or be covered by a window409. Alternatively, the GRIN lens 405 or other optical component canreside in the optically transmissive channel 407 of the acousticsubstrate 402, as seen in FIGS. 4 g to 4 l. The sleeve of insulatingmaterial 406 mentioned above can surround the GRIN lens 405 or otheroptical component within the opening 407 as shown in FIG. 4 l in orderto provide either mechanical or electrical insulation from the acousticsubstrate 402.

Referring to FIG. 4 f an optical spacer 433 is located between thedistal end of the optical fiber 403 and GRIN lens 405. The opticalspacer element 433 may comprise an optically transparent medium, such asno-clad fiber, glass, plastic, a gas-filled gap or a fluid-filled gap.The use of an optical spacer element 433 may help reduce the requiredprecision for the alignment and sizes of optical components in order toachieve a desired focal length.

Alternatively, as seen in FIG. 4 g, the path length of the prism 397 ormirror can act as all or a portion of the optical spacer 433 in betweenthe distal end of the optical fiber and the lens 405. The advantage ofusing the distance that light must travel through the mirror or prism397 as a substitute for a portion of a functional optical spacer is thatthe focusing element (e.g. the GRIN lens 405 or other lens) is closer tothe region being imaged, thus improving the effective working distanceof the optical imaging system. In some situations, the lens 405 can beoffset from either edge of the optically transmissive channel to achievethe desired depth of focus, as in FIG. 4 h.

In other embodiments, it may be helpful to have one or more opticalelements of the optical path extend beyond the outer surface of theacoustic transducer, such as element 434 as in FIG. 4 i, in order toachieve the desired performance of the optical imaging technique. Thisis particularly important when the acoustic transducer 402 is quite thin(such as a for very high ultrasound frequencies) or when the effectiveworking distance of the optical imaging technique is longer than can beaccommodated by having all the optical components reside below theemitting surface of the acoustic transducer.

It is also important to realize that the optical circuit can be distantfrom the surface of the acoustic transducer 402. By way of example, asseen in the embodiment shown in FIG. 4 j, it may be desirable to havesome backing material 435 interposed between the fiber optic 403 orother optical components proximal to the deflecting mirror or prism 397and the acoustic transducer 402 to minimize back-reflections from theoptical components.

The direction of propagation of the acoustic and optical imaging energycan be in a direction other than perpendicular to the longitudinal axisof the imaging probe. In fact, a slight angular offset of a few degreesis desired to minimize reflections back from the sheath that surroundsthe probe. FIG. 4 k shows an embodiment of a probe that combines opticaland acoustic imaging means aligned at an angle other than normal to thelongitudinal axis of the probe.

The embodiment of the probe 500 shown in FIG. 5 a is structurallyconfigured such that both acoustic and optical imaging sensors can becombined for viewing without components such as the mirror 404 of FIG. 4b or prism 397 of FIG. 4 c. The head section of probe 500 containingpiezoelectric material 402 for the acoustic sensor and the conductivelayers 401 on either side of the piezoelectric material 402 is alignedalong the longitudinal axis of the fiber optic 403 and the probe isconfigured so that both acoustic and optical signals are emitted axiallyrelative to the fiber axis, not perpendicular as in FIG. 4 a.

The embodiment shown in FIG. 5 b is analogous to the embodiment shown inFIGS. 4 b and 4 c. FIG. 5 c is analogous to the embodiment shown in FIG.4 d. The embodiment shown in FIG. 5 d is analogous to the embodimentshown in FIG. 4 e. The embodiment shown in FIG. 5 e is analogous to theembodiments shown in FIGS. 4 f and 4 g. The embodiment shown in FIG. 5 fis analogous to the embodiment shown in FIG. 4 i.

FIG. 6 a shows the geometry of an emitting surface of a squaretransducer 402. It should be noted that the geometry of the emittingsurfaces of the acoustic transducers 402 are not limited to being insquare in shape and may be any of several shapes, such as rectangular,circular, ellipsoid, and any other desirable shape. FIG. 6 b shows asquare transducer with the hole 407 in the center, while FIG. 6 c showsa square transducer with a glass rod 501 in the hole 407.

Results of a simulated beam profile using acoustic beam simulationsoftware are shown in FIGS. 6 d through 6 f, corresponding to thetransducer geometries in FIG. 6 a through 6 c respectively. As can beseen, there is considerable similarity in the beam profiles of thevarious configurations, providing evidence that ultrasound transducersadapted to allow a channel for optical transmission are capable ofproducing an acceptable ultrasound beam profile suitable for imagingpurposes.

A simpler method for aligning the optical and acoustic imaging meanswould be to place the fiber optic adjacent to the surface of theacoustic transducer 402 without going through the transducer 402 itself.FIG. 7A shows an imaging probe 510 comprised of an acoustic transducer402 with the distal end of an optical imaging circuit 428 placed on topof the acoustic transducer 402. The distal end portion of the opticalimaging circuit 428 comprises the distal end of fiber 403 and anyoptical components, such as an optical spacer 433, a lens, such as aGRIN lens 405, mirror 404 or prism 397, that enable emission orcollection of optical imaging energy. The distal end of an opticalimaging circuit 428 can be affixed directly to the acoustic transducer402 or supported by a support next to the acoustic transducer 402. Thedistal end of optical imaging circuit 428 would affect acoustic signalsgenerated and/or received by the acoustic transducer 402 as it liesdirectly in the path of a portion of the acoustic beam emitted bytransducer 402. However, a significant portion of the energy of theacoustic beam would not travel through the optical imaging means 403 andtherefore would remain relatively unaffected.

Furthermore, the signal processing means preferably includes signalsubtraction methods for discarding the portion of the signal thatrepresents the early time portion of an echo signal to cancelreflections from interfaces close to the acoustic transducer's surface.

FIG. 7 b shows a perspective view of imaging probe 512 which is amodification of the system in FIG. 7 a where the distal end of opticalimaging circuit 428 is recessed into the surface of the transducer 402thus rendering the recessed portion of the transducer non-functional, sothat acoustic beams transmitted or sensed by the acoustic transducer 402do not substantially propagate through the overlying imaging fiber 403.A top view of this embodiment is shown in FIG. 7 c. The portion of thetransducer 402 rendered non-functional can be rendered non-functional byeither removing the portion of the transducer 402 that lies underneaththe distal end of optical imaging circuit 428 as shown in FIG. 7 b, orby electrically isolating the portion of the electrode underneath theoptical imaging means. Removal may be done by several methods, includingthe use of a dicing saw to cut a channel through the transducer 402.Furthermore, removal of a channel makes it possible to considerrecessing the distal portion of the optical imaging means within achannel.

FIG. 7 c shows a top view of the emitting/receiving surface of the probe510 shown in FIG. 7 b surrounding the distal end of optical imagingcircuit 428.

FIG. 7 d shows an imaging probe 516 that employs a composite transducerfor the acoustic imaging means. In this case the composite transducer isa transducer comprising more than one signal generating element, orpillars 520. The composite transducer in FIG. 7 d comprises four pillars520. The channels 522 in between the pillars 520 leave a channel 522 forone or more distal ends of optical imaging circuit 428 to be placedwithin the confines of the composite acoustic transducer. The distal endof an optical imaging circuit 428 need not necessarily be recessedwithin channels 522, and can alternatively rest on or above the surfaceof the acoustic transducer 402. Conducting connections 400 between theupper conducting surfaces of the pillars 520 allows for the pillars tobe simultaneously activated. The channels 522 can be filled with afiller material, such as a polymer or epoxy, to increase the mechanicalstability of the composite transducer, or to help affix the opticalimaging means in place.

FIG. 7 e shows a top view of the imaging probe 516 with the distal endof the optical imaging circuit 428 placed within the center of thepillars 520. Any of the implementations for the distal portion of theoptical imaging circuit 428 (e.g. any combination of fiber optics,spacers, GRIN lenses, Ball lenses, air gaps, transparent windows), suchas those shown in FIG. 4, can be used in the implementations describedin FIGS. 7 a to 7 e.

As part of most mechanical scanning mechanisms for imaging, there is apredominant motion associated with the scanning mechanism that definesthe geometric path through which the imaging beam will sweep. Forexample, in an imaging system that uses a rotary motion to scan aregion, there will typically be a circular or conical surface, throughwhich the imaging beam sweeps, with the circular or conical surfacebeing centered approximately on the axis of rotation, as occurs incurrent implementations of mechanical scanning intravascular ultrasound.The predominant motion in this case is the rotational motion.

Alternatively, if the imaging emitter/receiver is translated along thelongitudinal axis, then the imaging beam will sweep through a planarsurface and the plane defined by that surface will include the axis oftranslation. This predominant motion in this case is a longitudinaltranslation.

If the imaging emitter/receiver is simultaneously rotated around alongitudinal axis of a probe and translated along a path that isgenerally parallel to the longitudinal axis of the probe, then theimaging beam will sweep through a surface defined by a helicoidgeometry.

It is possible to generate co-registered images with good precision frommultiple acoustic and/optical imaging means without having to have thetwo or more imaging beams be simultaneously collinear. This can beaccomplished by having one or more imaging beams follow the path of aleading beam. Software or electronic circuitry can use knowledge of thespeed and direction of the scanning mechanism's motions over time tothen register the images generated from one of the imaging means ontoone another.

For example, if the path of one imaging beam closely follows the path ofanother imaging beam (the leading beam) in a short time period, then itis possible to assume that the region scanned by the two means issimilar enough to accurately co-register the two images with each other.The accuracy of the registration between the two images can be affectedby the time delay in which the second beam follows the first beam. Ifthe time delay is relatively small, then inaccuracies in theco-registration of the two images that could potentially develop in thattime period are likely to be minimal. Such inaccuracies might includethose caused by tissue motion (such as that induced by cardiac orrespiratory motion), unintentional probe motion, physiologic changessuch as blood flow and imprecision in the fidelity of the scanningmechanism. The time delay (which itself can vary over time) can be usedfor the process of registering the different images.

FIG. 8 a shows an example of an imaging assembly 530 that contains bothan acoustic imaging means and an optical imaging means. The predominantscanning motion is a rotational motion around a longitudinal axis thatlies along the length of the imaging probe. As illustrated, the acousticimaging beam 532 and optical imaging beam 534 sweep through a path thatis circular in nature. If the imaging beams are not aligned normal tothe longitudinal axis, but rather at an angle other than 90 degrees fromthe longitudinal axis, than the path through which the imaging beamssweep will be conical in nature. If a longitudinal translation were tobe applied in combination with the rotary motion, the two beams wouldfollow a roughly helicoid path.

It will be understood that the in all embodiments disclosed herein theimaging assembly may be translationally movable within the hollow shaftand may emit anywhere along its length and is not restricted to thedistal end of the hollow shaft.

FIG. 8 b shows a side view of the combined imaging probe 530 where theacoustic beam 532 travels in one direction (upwards in the diagram) andthe optical imaging beam 534 travels out of the page (towards thereader). In this case, the optical beam 534 and acoustic beam 532 at anyinstant are oriented 90 degrees apart from each other.

FIGS. 8 c through 8 e represent a time series of the rotational motionof the imaging probe 530 as it would appear from the distal end of theimaging probe. In this example, the optical imaging beam 534 leads theacoustic imaging beam 532 by 90 degrees of rotation. At a constant framerate of 30 frames per second, the time delay that it would take for thetrailing beam to become collinear with a prior position of the leadingbeam would under 9 milliseconds, which is a short period of time withrespect to artifacts that might occur due to cardiac motion experiencedby an intravascular catheter.

Given the importance of miniaturizing the space occupied by componentsand assemblies in minimally invasive imaging means, it may be desirableto recess some of the components. For example, as seen in FIG. 9 a, animaging probe 540 has been configured to recess the distal end of theoptical imaging circuit 428 into the backing 435 of the acoustictransducer 402. Recessing may not only accomplish efficiency of spaceuse, but it may also provide a method of fixing the distal end of anoptical imaging circuit 428 to the acoustic transducer 402.

The purpose of the backing material 435 on the acoustic transducer 402is to attenuate signals generated from the back surface of thepiezoelectric 402 so that an image is not formed by the energy that isemitted from the back surface of acoustic transducer 402 on which theoptical emitter/receiver 403 is located, but rather only from theprimary emitting surface for acoustic signals (top surface) of thetransducer 402. Recessing an optical or other component in the backingmaterial 435 may potentially cause the optical or other component toreflect signals back to the acoustic transducer 402 that wouldpotentially create imaging artifacts.

FIG. 9 b shows a deflecting surface 544 in which the opticalemitter/receiver 403 is cradled that acts to deflect acoustic energythat might otherwise reach the optical emitter/receiver 403 and deflectsthat energy laterally (substantially parallel to the surface of theacoustic transducer 402) to minimize the amount of energy that isreflected back towards the transducer 402. This deflecting surface 544may be made of a hard substance such as glass or steel.

FIG. 9 c shows an implementation where the distal end of an opticalimaging circuit 428 itself has a surface 545 that substantially deflectsacoustic energy laterally without the need of an additional deflectingmaterial as seen in FIG. 9 b.

For embodiments of imaging probes where the imaging beams scan as aresult of rotational motion, it is not necessary that the rotationalvelocity be a constant or even remains in the same direction. It ispossible to have a reciprocating motion where the imaging assemblyrotates in one direction and then stops and rotates in the oppositedirection. In this situation, the leading and trailing beams swap roleswith each other.

For example, in FIG. 10 a, the acoustic beam 532 initially follows theoptical beam 534 as the imaging assembly rotates in a counter clockwisedirection. The acoustic beam 532 continues to follow the sweep path ofthe optical beam 534 as shown in FIG. 10 b until the rotational velocityof the imaging probe reaches zero, (as in FIG. 10 c). Once the directionof rotation changes to the opposite direction, the acoustic beam 532becomes the leading beam and the optical beam follows (as in FIGS. 10 dand 10 e). The motion can change direction as many times as desired witha concomitant change in the definition of the leading and trailingsensor beams.

FIG. 11 shows an imaging probe 540 where the predominant motion is alongitudinal motion back and forth along arrow 541 where the surfaceswept the optical beam 534 and the acoustic beam 532 are two co-planarrectangles. As the imaging assembly is translated proximally (to theleft in FIG. 11) the optical imaging beam 534 leads the acoustic imagingbeam 532. The opposite is true for distal translation (to the right inFIG. 11). The longitudinal motion can be reciprocated as well.

With either longitudinal or rotational predominant motions, it isunderstood that additional motions can be combined with the predominantmotion. For example, a slow translation (such as 10 mm/s or less, andtypically 1 mm/s or less) can be added to a rapid rotational scanningmotion (such as 360 degrees per second or more and typically 3600degrees per second or more) in order to acquire 2D cross-sectionalimages at different longitudinal positions.

Similarly, a slow rotational motion (e.g. less than 360 degrees persecond and typically less than 30 degrees per second) can be added to asequence of rapidly reciprocating longitudinal motions (averaging over0.1 mm/s and more typically more than 1 mm/s) to create a series oflongitudinal images acquired at different orientations around thelongitudinal axis of the imaging probe. The alignment of the variousimaging elements at the distal end is configured such that the one ofthe imaging beams will follow the other during the predominant motion,but the ability to accurately register the images on top of each otherwould not be significantly affected by the addition of a relative slowsecondary motion. While absolute numbers for slow and rapid motions inthe rotational and translation motions are provided above, it is therelative magnitude of these motions that is more important.

Collinear alignment of the optical and acoustic beams (as shown in theembodiments shown from FIGS. 4 a to 5 f) provide very accurateregistration of the optical and acoustic images. An alternativeembodiment of the probe is configured to have the optical and acousticbeams substantially overlap each other by angling either the opticalimaging emitters/receivers towards the path of the acoustic beam or byangling the acoustic imaging emitter towards the path of the opticalimaging beam. FIG. 12 shows such an embodiment of an imaging probe 546where the distal end of an optical imaging circuit 428 is configuredsuch that the optical imaging beam 534 is angled towards the acousticimaging beam 532 and vice versa. This provides a simpler method ofconstruction than aligning the optical and imaging beams as seen inFIGS. 4 a to 5 f, but allows the two imaging means to provide what maybe a reasonably precise overlap over a portion of the two imaging beams.In particular, embodiments whereby the beams are aligned such that theyoverlap over a substantial portion of their focal ranges would beuseful.

FIG. 13 shows an embodiment of the imaging probe 550 configured to imagesimultaneously in the same general orientation and from the same generalorigin with both acoustic and optical means. At least one fiber optic410 and one electrical conduit 411, such as a pair of coaxialconductors, reside within the imaging conduit 560 and travel to theimaging assembly 562. The imaging assembly 562 comprises an acoustictransducer 412 configured for imaging in a substantially side-viewingdirection indicated by arrow 420. The imaging assembly 562 also includesdistal end of an optical imaging circuit 564 configured for imaging in asubstantially side-viewing direction indicated by arrow 421.

The acoustic transducer 412 and distal end of an optical imaging circuit564 are configured such that they allow imaging in two or more separatedirections at any instant within the same cross-sectional plane that issubstantially perpendicular to the axis 423 around which the imagingassembly 562 rotates. Thus, assuming minimal translation of the imagingassembly 562 while the imaging assembly is rotated, the imaging datacollected by the optical emitters/receivers 564 can be co-registeredwith the imaging data collected by the acoustic transducer 412. Forexample, if the acoustic and optical means are configured to image indirections that are 180 degrees opposite of each other around thelongitudinal axis, as shown in FIG. 13, then the region imaged by theacoustic transducer 412 at one point in time will be substantially thesame region that is imaged by the distal end of an optical imagingcircuit 564 after the imaging assembly 562 has been rotated by half arevolution. Similarly, if the imaging beams 420 and 421 have a similarangle from the longitudinal axis other than 180 degrees, they will bothsweep through paths of substantially coincident cones, and can thereforebe co-registered.

The embodiment of the probe 570 shown in FIGS. 14 a and 14 b isconfigured such that both IVUS and OCT can be combined for forwardviewing with a deformable component. At least one fiber optic 410 andone electrical conduit 411, such as a pair of coaxial conductors residewithin the imaging conduit 578 and travels to the imaging assembly 572.The acoustic transducer 412 is configured for imaging in a substantiallyforward-looking direction indicated by arrow 413. A distal end of anoptical imaging circuit 574 is configured for imaging in a substantiallyforward-looking direction indicated by arrow 414.

The distal end of an optical imaging circuit 574 typically comprises adistal end of a fiber optic 410 combined with a lens 415, such as a GRINlens and an optional spacer (not shown). The imaging conduit 578comprises an artificial muscle actuator that has the property of beingable to deform upon the application of an electrical charge. FIG. 14 billustrates how the imaging angle would be changed if an artificialmuscle actuator achieved a deformation while FIG. 14 a shows the shapeof the probe without application of a voltage to actuator.

Embodiments of the present imaging probe may be configured to make useof a deflector to allow for a larger transducer to be used within theimaging probe. Alternatively, the deflector may be pivotable and coupledto a pivoting mechanism to enable an additional degree of freedom in thescanning mechanism. For example, the scanning mechanism may facilitate2D imaging, or may augment a 2D imaging system into a 3D imaging system.Alternatively, the deflector may be translated along the longitudinalaxis in order to change the focal depth of the imaging system.

FIG. 15 a illustrates an embodiment of an imaging assembly 590 thatcomprises a deflector 592 used to deflect optical and/or acousticimaging energy into a generally radial direction. The deflector 592 ismade of one or more reflective materials. Optically reflective materialsinclude polished or sputtered metals, such as stainless steel, gold,silver and platinum.

Acoustically reflective materials include stainless steel and othermetals, quartz and other crystals, glass and hard polymers. FIG. 15 bshows another embodiment of an imaging assembly 600 which comprises adeflector 602 that pivots around a pivot point 604 and thus allows theangle between the imaging beam and the longitudinal axis of the imagingprobe to vary. The imaging assembly 600 may be configured so thatdeflector 602 can change position by being coupled to a variety ofmechanisms, including mechanisms which utilize centripetal motion,magnetic forces, cable mechanisms, rheologic forces, piezoelectricdrivers, miniaturized motors and others.

FIG. 15 c illustrates an embodiment of the arrangement in FIG. 15 bwherein a cantilever 901 mounted on a cantilever mount 902 and thedeflector's range of motion is limited by a minimum stop 82 and amaximum stop 80. This embodiment has the property of having the imagingangle change as a result of changes in the rotational motion of theimaging assembly around the longitudinal axis of the probe. At rest orlow rotational speeds, the cantilever wire forces the deflector 602around its pivot point such that it comes into contact with stop 80. Athigher rotational speeds, centripetal acceleration causes the deflector604 to pivot away from stop 80. As centripetal acceleration continues tooverpower the restoring force exerted by cantilever 901 on deflector602, the deflector eventually comes into contact with stop 82. In suchan embodiment, an imaging assembly 600 with a 3D scanning mechanism isimplemented.

FIG. 16 a illustrates an embodiment of the distal portion of an imagingprobe 100 capable of both acoustic and optical imaging in a generallyforward-looking direction. FIG. 16 a shows an embodiment of a distal end29 of an imaging probe containing an imaging assembly 30 that includes atiltable component 70 where the tiltable component is a disc mounted ona pivoting mechanism such as a pin 72 that extends through the disc 70.The pivoting mechanism 72 defines the tilting axis of the tiltable disc70. When the imaging assembly 30 is at rest, the disc 70 will remain inan arbitrary starting position. However, as the imaging assembly 30rotates, the disc 70 will align itself such that the normal of theplanes defined by the faces of the disc 70 are substantially parallelwith the longitudinal axis 75. The disc 70 has two preferredorientations when the imaging assembly 30 is rotated, that are separatedby a rotation around the tilting axis of 180 degrees.

For the purposes of this description, the tilt angle will be referred toas the angle between the longitudinal axis 75 and an imaginary axisthrough the tiltable component 70 that is parallel to the longitudinal75 axis when the tiltable component 70 is in one of its preferredorientations. By way of example, when the tiltable component 70 is in apreferred orientation, the tilt angle is approximately zero. If thetiltable component 70 is tilted away from its preferred orientation byan external force, such as gravity, magnetic forces, electrostaticforces, friction with another moving part or fluid, compressive forces,normal forces or any other source of incompletely opposed torque on thetiltable component 70 around the tilt axis, the tilt angle willincrease.

One or more mechanisms may be included in the imaging assembly 30 thattends to cause the tiltable component 70 to have its tilting angleincrease. For the purposes of this invention, such a mechanism isreferred to as a restoring mechanism. A torsion spring 76 (as shown inFIGS. 16 a and 16 c), a cantilever or a compression spring can be usedas a restoring mechanism, where one end of the spring 76 is mechanicallyin contact with tiltable component 70 and the other end is mechanicallyin contact with another part of the imaging probe 100, such as the bodyof the imaging assembly 30.

Alternatively, magnetic, electrostatic, hydraulic or other mechanismsthat apply a torque on the tiltable component around the tilting axiscould be applied. Other examples of mechanisms that could be used toprovide a restoring force include tension from an elastomer (such asrubber, polyurethane, silicone, fluoroelastomers, thermoplastics andmany others) or by use of a cantilever spring or foil, such as springsor foils made of platinum, nitinol, steel or other suitable materials.In very small embodiments of the imaging device, where intermolecularforces such as electrostatic forces and Van der Waals forces betweencomponents in the imaging assembly may become quite significant evenwithout the application of an external voltage. Therefore, the innateintermolecular forces between the tiltable component and structuresclose to the tiltable component, such as the stops 80 and 82 describedbelow, may be sufficient to provide a net restoring force. For example,a stop comprising a surface made of PVC, nylon or LDPE could providesufficient attraction between the tiltable component and the stop.

One or more stops 80 and 82 may limit the range of the tilt angle of thetiltable component 70. For example, a post or lip 80 can extend from theshell 84 of the imaging assembly 30 as a stop to prevent the tiltingcomponent from further changing its tilt angle while it makes contactwith the stop 80. Therefore, a stop can be used to limit the tilt anglefrom exceeding a maximum value determined by the position of the stop.In many embodiments, this maximum tilt angle is the tilt angle that isachieved when the imaging assembly 30 is at rest and at low rotationalspeeds.

An additional or alternative stop 82 can be included to create a minimumtilt angle that the tiltable component will achieve at rotational speedsin the upper end of the operating range. Indeed, there are manysituations in which there is no significant benefit in allowing the tiltangle to reach zero, as will become apparent in the followingdescriptions of specific embodiments. FIG. 16 c shows the tiltablecomponent hitting the second stop to limit its range of motion at higherrotational speeds of the imaging assembly.

The imaging assembly may include both optical emitters and associatedoptics and ultrasound transducers. The ultrasound transducer 88 ismounted at the end of small coaxial cable 89 and lens 92 and mirror 94are mounted at the end of a fiber optic cable 96 in the imaging assembly30 in FIGS. 16 a to 16 d with the optical and ultrasonic emittersconfigured to focus imaging energy onto the tiltable component 70. Theultrasound transducer 88 and optical emitter can direct imaging energytowards the tiltable component 70. Alternatively, one of the embodimentsthat enables collinear optical and acoustic imaging, as seen in FIGS. 4a through 4 k or FIGS. 5 a through 5 f can direct imaging energy towardsthe tiltable component 70.

The imaging energy is then deflected by an energy-deflecting componentmounted on the tiltable component 70. For ultrasound imaging, theenergy-deflecting component (the tiltable component 70) may comprise anacoustically reflective surface, such as a solid metal surface (e.g.stainless steel) or crystalline surface, such as quartz crystal orglass. For optical imaging, the energy deflecting component (tiltablecomponent 70) can comprise an optically reflective surface such as amirror surface made from polished metal, metallized polymer such asmetallized biaxially oriented polyethlylene terephthalate (Mylar),sputtered or electrochemically deposited metal or metal foil. Metalscommonly used to make mirrors include aluminum, silver, steel, gold orchrome.

Alternatively, the energy-deflecting component could be made of atransparent refractive material, such as glass, clear polymers, and manyothers, and deflect the imaging energy in a manner similar to a prism.Preferably, the emitter and/or receiver is mounted on a component of theimaging assembly that rotates with the imaging assembly. However, it isalso possible that the emitter and/or receiver is mounted on a componentof the imaging probe that does not rotate with the imaging assemblywhile the energy deflecting mechanism within the imaging assembly doesrotate. This could be achieved by mounting the emitter and/or receiveron an external sheath for example, or by having the imaging assemblydivided into two or more sub-assemblies, one of which rotates andincludes the tiltable component.

For ultrasound and optical coherence tomography, the ability to adjustthe angle of propagation of the emitted and/or received imaging energy,when combined with the rotational motion of the imaging assembly, allowsa 3D volume to be scanned. For angioscopy and infrared imaging, theability to adjust the angle of propagation of the emitted and/orreceived imaging energy, when combined with the rotational motion of theimaging assembly, allows an image to be produced using a single fiberoptic rather than requiring a bundle of fibers. Such an improvement canresult in greater flexibility and/or miniaturization of the imagingdevice.

Further details of various scanning mechanisms that may be used in theimaging probe disclosed herein are disclosed in U.S. patent applicationSer. No. 12/010,206 entitled SCANNING MECHANISMS FOR IMAGING PROBE,filed Jan. 22, 2008, now U.S. Pat. No. 8,214,210, which his incorporatedherein by reference in its entirety.

In the case where the energy-deflecting component comprises a reflectivesurface it is not necessary that the reflective surface be planar. Forexample, in the case of acoustic imaging, it may be advantageous for anacoustically reflective surface to have a contour to it, such as aparabolic or spheroid contour, so that the acoustic beam can be focusedby the reflective surface and improve lateral resolution of the acousticimaging system as a result. Furthermore, in the case where the tiltingcomponent is used to deflect both acoustic and optical energy usingreflection, the acoustic reflector need not be the same surface thatreflects the optical energy.

For example, while it might be advantageous to have a contour such as aparabolic contour for the acoustically reflective surface, it may bepreferable to have a planar surface for the redirection of the opticalimaging energy. This can be accomplished by having an acousticallyreflective surface such as a stainless steel disc with one of its facescontoured to have a parabolic shape to it as in FIGS. 17 a through 17 dwhich show a tiltable deflecting component that has an opticallyreflective surface that is distinct from the acoustically reflectivesurface.

FIG. 17 a is a perspective drawing of a deflector that has holes on itsside for receiving pins on which the deflector can pivot within animaging assembly. FIG. 17 b shows a cross-section through the deflectornear the center of the deflector. The holes for receiving pins 465 areseen. The top layer is a flat, optically reflective layer 461. Under theoptically reflective layer 461 is a generally acoustically transparentlayer 462, which lies between the optically reflective layer 461 and anacoustically reflective substrate 463. FIGS. 17 c and 17 d showcross-sectional images of such a deflector at different points away fromthe center of the disc.

Such a deflector can be constructed by taking a disc of an acousticallyreflective material such as stainless steel and drilling the necessaryholes or indentations so that the deflector can eventually be mountedinto an imaging assembly. A parabolic or spheroid indentation can bemade into one face of the disc. The indented surface can then be filledwith an acoustically transparent medium, such as polymethylpentene(TPX). A thin layer of gold, silver or chrome can be sputter depositedonto the exposed planar polymer surface to act as an opticallyreflective surface. Such a layer may be on the order of 300 Angstroms to20,000 Angstroms such that it is thin enough that its mechanicalproperties to allow acoustic energy to transmit through it, whilesimultaneously providing an optically reflective surface.

The result of such a fabrication process is to create a layeredreflector that reflects acoustic energy from the contoured surface toachieve the desired focusing effect, while the optical energy isreflected from a planar surface. It is a further advantage of thisconstruct that the optical and acoustic imaging can occur in aconfiguration where the optical and acoustic imaging energy travelsthrough the same general space, facilitating co-registration of opticaland acoustic images and minimizing the amount of space required withinthe imaging assembly to accommodate more than one modality of imaging.

In some embodiments, such as the assembly shown in FIGS. 16 a and 16 c,it may be helpful to use one of the imaging modalities solely to measurea parameter useful for the reconstruction of 2D and 3D images. Forexample, in the case of a volumetric imaging probe that uses adeflectable component, it may be desirable to use OCT to accuratelymeasure the tilt angle of the deflectable component. Thus, an ultrasoundimage could be generated with knowledge of the tilt angle derived fromOCT data, such as the tilt angle of tiltable component 70 in FIG. 16 awithout necessarily using the OCT data to generate corresponding OCTimages of the region outside of the imaging probe.

In some embodiments, it will be desirable to have more than one methodfor optical imaging in an intravascular imaging system. For example, OCTand angioscopy may be a useful combination. FIG. 18 a shows anultrasound imaging transducer 402 with two (2) distal ends of opticalimaging circuits 428 through two (2) separate optically transmissivechannels in the acoustic transducer. FIGS. 18 b and 18 c show anacoustic imaging transducer with two (2) distal ends of optical imagingcircuits 428 arranged in a manner such that they are aligned along thepredominant rotary motion of the imaging assembly. These are examples ofusing more than one optical imaging emitter/receiver at the distal endof the imaging probe. If the imaging probe uses extensive rotary motionaround its longitudinal axis as part of the scanning mechanism, suchembodiments may require the use of a multi-channel optical rotary joint.

Alternatively, the optical imaging light sources and/or detectors forsome of the imaging systems may be mounted on the rotary portion of theimaging probe and be coupled to the imaging system using electrical sliprings or wireless communication. A battery may optionally be used as asource of electrical energy on the rotary portion of the probe oradapter to minimize the number of slip rings required. Illuminatingsources and photodetectors can be placed at the proximal end of theimaging probe and may be configured such that they rotate around thelongitudinal axis of the probe with the rest of the imaging conduit 34so that further optical couplers are not required between the imagingprobe and the adapter. This is done because the complexity of rotaryoptical joints increases substantially if more than one fiber isinvolved to connect the probe to the rest of the system.

If the imaging probe uses only reciprocal rotary motion over a shortrange of angles (such as less then two full revolutions), or no rotarymotion at all, then the use of an optical rotary joint is not necessary,simplifying the task of coupling the optical elements of the imagingprobe to the image processing and display hardware.

The imaging probe may include a motion detector for detecting movementof the movable member (tiltable or bendable members) relative to aremainder of the imaging assembly. The motion detector may be based onany of optical coherence based detection means, reflection intensitydetection means, and a strain gauge based detection means.

The pivotally mountable members may be pivotally mounted on a lowfriction pivot mechanism. The restoring mechanism is provided by any oneor combination of a spring and a magnetic/electromagnetic assembly asdiscussed above. The restoring mechanism may also include a surfaceexhibiting electrostatic properties which interact with the movablemember. It will be understood that the hollow shaft may be an externalcatheter sheath which may have memory properties.

All embodiments of the imaging probe disclosed herein may be fitted toexisting control and image processing system and display systems towhich the probe is connectable. The processing and display system wouldbe configured to process the received energy signals and produce imagesof interior surfaces or adjacent structures of said bodily lumens andcavities or exterior surfaces or adjacent structures of a body.

In another embodiment, it is possible to using the same optical imagingemitter/receiver at the distal end of the imaging probe and use opticalrouting circuitry such as switches, multiplexers, demultiplexers,prisms, diffraction gratings, couplers and/or circulators to use thesame fiber and distal optical components for more than one imagingmodality. FIG. 19 shows a schematic of a system where there are two (2)optical imaging systems 211 that are coupled to the same optical imagingwaveguide 212 via optical routing circuitry (comprising one or more ofthe components listed above). The waveguide may be coupled to theimaging probe via an optical rotary joint 213 if the image probe 12requires a large range of rotary motion as part of its scanningmechanism. The distal end of optical imaging circuit 428 may compriseany of the combinations of optical fiber, spacers, mirrors, prisms, balllenses, GRIN lenses, air gaps and transparent windows mentionedelsewhere in the present invention to enable optical imaging. While manyoptical imaging elements, such as the waveguide and lenses, are designedto operate optimally for particular ranges of wavelengths (e.g. infraredvs visible spectrum), the performance of a fiber optic or other opticalcomponent designed for one range is often still adequate to provideinformation using light in the other spectrum.

Therefore, imaging using more than one range of wavelengths can occursimultaneously. Alternatively, the imaging waveguide can be used atdifferent time intervals for different imaging modalities by means ofoptical switches, multiplexers and demultiplexers within the opticalrouting circuitry 210, or by simply timing the use of the opticalwaveguide at different time intervals for different imaging modalities.

While a fiber optic would be a preferred optical waveguide 212 for mostembodiments, it may be desirable to use an alternative form of opticalwaveguide that is potentially more space efficient than an opticalfiber. For example, a thin optical channel, on the order of 3 to 500microns in maximal diameter and preferably on the order of 4 to 125microns can be formed in a catheter at the time of extrusion. A fluidmedium with a high index of refraction can be introduced into theoptical channel, such as by means of injection. Such a fluid medium mayinclude an epoxy or adhesive specifically designed for opticalcomponents.

The fluid medium may also be curable, such as in the case of UV curableadhesives. The creation of an optically transparent channel filled witha material with a high index of refraction surrounded by the extrudedcatheter material with a lower index of refraction would essentiallyreplicate the functionality of including a fiber optic, but may allowfor slightly more efficient use of space in the catheter by notrequiring a separate cladding layer. The optimal use of space in acatheter is often important given their minimally invasive nature andthe limited space available in the regions in which these catheters aredeployed.

Yet another mode of operation for the present invention is the use of atransducer that combines acoustic transduction with an opticaltransducer where the transmitted energy is of one form and the receivedenergy is of another. For example, photoacoustic imaging comprisesdelivery of light-based energy to an imaged region. The photons interactwith the imaged region and create acoustic energy as part of theirinteraction with the medium in which they propagate. This acousticenergy is often in the form of ultrasound waves, and can be detected byan ultrasound transducer. It should be apparent that the use of anoptical emitter aligned and in combination with an acoustic receiverwould be a good configuration to enable photoacoustic imaging. Anultrasound transducer with an opening for optical imaging or that allowssubstantial overlap in the acoustic and optical imaging regions, such asthose shown in FIGS. 4 a through 4 k, 5 a through 5F or FIG. 12, wouldenable photoacoustic imaging.

Similarly, sonoluminescent imaging comprises delivery ofultrasound-based energy to an imaged region (Daniels and Price,Ultrasound in Medicine and Biology 1991:17(3):297-308). The acousticenergy interacts with the imaged region and creates photons as part ofits interaction with the medium in which it propagates. Some of thesephotons are directed back toward the source of the acoustic energy. Itshould be apparent that the use of an ultrasound transducer aligned incombination with an optical receiver would be a good configuration toenable sonoluminescent imaging. Implementations of acoustic and opticalimaging elements where the imaging beams are collinear, or substantiallyoverlap, such as those shown in FIGS. 4 a through 4 k, 5 a through 5 for FIG. 12, would enable sonoluminescent imaging.

Referring to FIG. 1 again, imaging probe 12 (which may include any ofthe embodiments of the acoustic and optical sensors discussed herein)and its components may be of several dimensions and properties dependingon the anatomic location and purpose of use for the imaging that isenabled by the imaging probe 12. For example, for the purposes of use inthe cardiovascular system, including the cardiac chambers, the imagingprobe 12 would preferably be elongate and flexible, with a lengthranging from 5 to 3000 mm, preferably with a length ranging from 300 mmto 1600 mm. The imaging conduit 34 and imaging assembly 30 may have amaximum cross-sectional dimension ranging from 200 microns to 10 mm,preferably ranging from 500 microns to 5 mm. An external sheath 48 maysurround both the imaging conduit 34 and imaging assembly 30. This wouldenable the imaging conduit 34 and imaging assembly 30 to rotate withinthe external sheath while mechanically isolating the rotational motionof these two components from the surrounding tissues.

In yet another example, the use of the imaging probe 10 in thegastrointestinal system would typically have the imaging probe 10 beingelongate and flexible, with a length ranging from 100 mm to 2000 mm andpreferably in the range of 300 mm to 1500 mm. The maximumcross-sectional dimension would typically range from 3 mm to 20 mm.

In yet another example, the use of the imaging probe 10 to image softtissue via percutaneous means would have the imaging probe with a rigidshaft. The external sheath would be replaced by a rigid hollow shaft,such as a stainless steel tube although many other polymers, metals andeven ceramics would be functionally suitable.

In yet another example, the use of the imaging probe 10 in theintraoperative neurosurgical setting would typically have the imagingprobe 10 being short and semi-flexible, with a length ranging from 50 mmto 200 mm. It would be preferable that the surgeon can bend and shapethe probe during the procedure to provide optimal passage fromextra-cranial space towards the intracranial target being imaged. Themaximum cross-sectional dimension would range from 200 microns to 5 mmand preferably from 500 microns to 3 mm.

In yet another example, the use of the imaging probe 10 in theinterventional neurovascular setting would typically have the imagingprobe 10 being long and ultraflexible, with a length ranging from 200 mmto 4000 mm and preferably ranging from 1300 mm to 2000 mm. The maximumcross-sectional dimension would range from 200 microns to 5 mm andpreferably from 500 microns to 3 mm. The distal end of the probe wouldpreferably possess shape memory to enhance navigation through theneurovasculature.

Embodiments of the present invention can be used in conjunction with orincorporated into devices that are used for intervention, such as thoseused for cardiovascular intervention, such as an angioplasty balloon,atherectomy device, stent delivery system or localized drug deliverysystem. It can also be used in conjuction with or incorporated intodevices that facilitate biopsies, radio-frequency ablation, resection,cautery, localized brachytherapy, cryotherapy, laser ablation oracoustic ablation.

In particular, using the image scanning mechanism to direct higherpowers of optical or acoustic energy to a targeted region can facilitatethe use of the current device to enable laser or acoustic ablation oftissue. For example, while imaging a region of a blood vessel with anOCT or ultrasound embodiment of an imaging probe described in thepresent invention a region for the delivery of therapy can be selectedthrough a user interface. Then, powerful pulses of energy can bedelivered at times when the scanning mechanism is oriented to deliveryenergy in the desired direction. For example, pulses of laser energy canbe transmitted down the same fiber optic used for optical imaging, bedeflected by a deflecting component in those embodiments that include adeflecting component, and travel towards the targeted tissue for thedesired effect. The timing of the pulses of laser energy is coordinatedwith the scanning pattern realized by the imaging probe to direct theenergy towards the targeted region.

The opportunity to acquire accurately registered images of two or morehigh resolution imaging modalities provides significant information thatis likely to be more useful than available by a single imaging modality.Maschke et al describe the formation of a composite image whereby theinner portion of an intravascular image is composed of OCT imaginginformation while the outer portion of an intravascular image iscomposed of IVUS imaging information. This takes advantage of the higherresolution images acquired by OCT and the higher penetration of IVUS.However, the reliability of this superposition of IVUS and OCT images islimited by the inaccuracy of the registration in the IVUS and OCT imagesthat occurs using the arrangement of the IVUS and OCT imaging elementsas described by Maschke and are substantially overcome by many of theembodiments in the present invention.

Alternative presentations of combined IVUS and OCT images might includedividing the image into sectors, where alternating sectors are displayedusing alternating imaging means, as seen in FIG. 20 a. First image 231and second image 232, where the first and second images areco-registered with each other images and acquired by different means,can be used to form a combined image 234 where sectors 233 of the firstimage replace sectors of the second image. Optionally, the borders 235defining the sectors 233 can rotate over time around the center of theimage to provide a dynamic image for identifying features in both thefirst and second co-registered images. FIG. 20 b shows a timeprogression of the rotations of the borders 235 around the center of thecombined image 234.

Alternatively, the user can specify which portions they would like tohave as one image and which they would like to see as the other byidentifying closed contours 236 in the second image as seen in FIG. 21 aor by identifying a space 237 in between two closed contours in thesecond image, as seen in FIG. 21 b.

Alternatively, displaying the first image 231 and second image 232 atthe same position on the screen as separate layers and varying thetransparency of the layer in the foreground can effectively provide ameans for combining the images. Alternatively, the order of the layerscan be varied over time, such as by having the IVUS image in theforeground for one time interval and then transitioning to having theOCT image in the foreground for a subsequent time interval, as seen inFIG. 22.

It is an object of the present invention to be able to identify certainfeatures of interest in a first image 231 and transfer knowledge of thatfeature (such as its position, shape, signal properties or composition)to a second image 232 that is accurately co-registered with the firstimage 231. Geometric features include specific points, contours or 2Dregions in an image. As seen in FIG. 23 a, a user can identify a point238, contour or region in a first image 231 manually, through the userinterface of the imaging system (such as with a mouse or keyboard) andhave that geometric point 238 appear in a second image 232 co-registeredwith the first image 231 as in FIG. 23 b. The availability of one ormore other images that are accurately co-registered with the first imagemakes it possible to superimpose any or all of the geometric featuresfrom the first image to any of the other images.

By way of example, the user might identify the inner boundary of a bloodvessel or the trailing edge of a fibrous cap in an OCT image. FIG. 24 ashows the contour representing the inner border 241 identified in aschematic representation of an OCT image (the first image). Similarly,the outer boundary 242 of the vessel wall (usually defined by theexternal elastic lamina) can be identified in an IVUS image (the secondimage). The contours representing the inner boundary 241 of the bloodvessel or the trailing edge of the fibrous cap can then be superimposedonto the corresponding IVUS image. Similarly, the outer boundary 242 ofthe vessel wall (usually defined by the external elastic lamina) can beidentified in an IVUS image. The contour representing the outer boundaryas assessed in the IVUS image can be superimposed onto the OCT image.FIG. 24 b shows the inner and outer boundaries on both the first andsecond images.

While the inner boundary of the blood vessel is readily identified onmost IVUS images, the OCT generated contour would be more accurate inmost circumstances. Furthermore, OCT is thought to be much better foridentifying the fibrous cap of a plaque, in part due to its higherresolution. However, IVUS can see much further into most vasculartissues and can provide a better assessment of the outer vessel wall.

A geometric feature can include features observed in 3D data sets, suchas surfaces or volumes. A surface or volume observed in a 3D imagingdataset can be superimposed into another 3D imaging dataset if the twoimaging datasets are accurately registered.

The geometric features of interest need not be manually identified. Itis possible that features in an imaging dataset can be identified byautomated or semi-automated means to minimize user intervention. Forexample, there are several border detection methods cited in theliterature on IVUS (e.g. Klingensmith, IEEE Transactions on MedicalImaging, 2000; 19:652-662). Automated border detection methods analyzean image to identify a contour of some pre-determined significance.Semi-automated methods are similar, but require some user interventionto either provide a starting point for the border detection algorithm orto refine the results produced the algorithm.

Other feature detection algorithms can be conceived of to identifyfeatures other than a border. For example, a hyper-intense/bright regionin an ultrasound image followed a dark region in the same direction ofthe imaging beam is often referred to as “shadowing” and occurs mostcommonly when the area being imaged includes either calcium (such asfrom advanced atherosclerosis or malignant processes) or metal (such asfrom stents or other implants). Similarly, a highly intense region in anOCT image of a blood vessel, followed by a rapid but continuousattenuation of the signal acquired along the same imaging path issuggestive of necrotic material in the vessel wall. It is possible todetect such regions algorithmically and identify them in theirrespective images. Once such features are identified in their respectiveimages, their position and shape can be superimposed into other imagesthat are accurately co-registered.

In certain embodiments of the present invention, it will be desirable todo some adjustment to one or more of the images to further improve theco-registration. While many of the embodiments of the present inventionimprove the precision of acquiring imaging data with one or more imagingmethods, there may be some advantage to further adjusting the images toimprove the accuracy of the co-registration process. For example,ultrasound images are generated assuming a constant speed of soundthrough all tissues, while OCT assumes a constant speed of light throughall tissues.

In reality however, there are small changes in these speeds depending onthe composition of the tissue in which each of the imaging energiespropagate. Therefore, prior to completing the co-registration processfor one or more images, it may be desirable to morph or warp one or moreof the images by identifying certain features in the two or more imagesthat are to be co-registered and using those features to guide themorphing process. Any point, contour or other feature identified in allof the images to be co-registered can be used to drive the morphingprocess. An ultrasound image is most commonly formed by displaying agrayscale representation of the intensity of the ultrasound signalreflected back from the approximate anatomic location that correspondsto each pixel in the image. Similarly, an OCT image is most commonlyformed by displaying a grayscale representation of the intensity of thelight reflected back from the approximate anatomic location thatcorresponds to each pixel in the image.

Aside from the intensity information at each location in either anultrasound or OCT image, there are several other features fromultrasound or OCT images that can be very helpful for analysis derivedfrom combined imaging.

The display of an image derived from ultrasound signals based on afeature other than then intensity of a sample in the image is well knownin the art. Nair et al (Circulation 2002; 106(17):2200-2206 and U.S.Pat. No. 6,200,268) published results of an algorithm that measuresseveral parameters of an ultrasound signal in discrete regions of IVUSimages of blood vessels. Each region was also assigned a tissue categorybased on histological analysis of the vessel. The ultrasound derivedparameters and the histological classification of each region were inputinto a pattern recognition engine to generate an algorithm that issubsequently applied in an attempt to classify tissue in vivo based onits many ultrasound signal properties. Some of the properties used foranalysis include frequency domain parameters over a defined range offrequencies such as maximum power, frequency of maximum power, minimumpower, frequency of minimum power, slope, y-intercept, mid-band fit andintegrated backscatter. The image generated comprises a topographicalmap of the vessel cross-section and a discrete number of colors, witheach color representing a single tissue category. Wilson et aldemonstrated the use of measuring the frequency domain attenuation of anultrasound signal in regions of an IVUS images and overlaying a colormap of the attenuation slope onto the conventional IVUS image toidentify areas thought to correspond to specific pathological types.

Similarly, features of interest can be measured or identified in opticalimages in order to generate images other than intensity-based images.Parameters or other features that can be used to generate such imagesinclude attenuation, polarization sensitivity, detected edges,spectroscopic information and others.

As a result of the high degree of accuracy of co-registration enabled bythe present invention, it is possible to generate images based onfeatures or signal properties measured with more than one imagingmodality. For example, a composite image can be made using an innerborder 245 identified by OCT, an outer border 246 identified by IVUS anda color map of the most likely tissue components within the vessel wallusing a pattern recognition system that combines optical signalproperties with acoustic signal properties within focal regions of theimaging datasets to generate a composite image that will improve theability to identify important components within the vessel wall, such ascalcified, fibrous, atheromatous, thrombotic, metallic and non-diseasedregions.

FIG. 25 a shows a schematic representation of an inner border 245identified by OCT, an outer border 246 identified in the second image byIVUS and a region of interest 247 used for analysis of the OCT andultrasound signal properties. As shown in FIG. 25 b, the signalproperties 248 from the more than one modalities of imaging in theco-registered region of interest are used to generate an assessment ofthe composition of one or more pixels in the composite image thatcorrespond to the region of interest analyzed. The assessment may beformed by a pattern recognition system 249 trained using methods knownin the art. The geometric features 249 identified in the co-registeredimages are also optionally included in the composite image. The processof assessing the composition of a region of interest can be repeatedseveral times over for different regions of interest to generate acomposite image.

In addition, the software and image processing algorithms that enablessuch analysis of the combined imaging means need not be on theacquisition station. Once the imaging data is acquired, the imaging datacan be transferred to allow analysis to occur offline on a separate setof one or more processing units.

The combined IVUS/OCT scanning devices disclosed herein may include arotary encoder. Further details of optical encoders which may used withthe combined IVUS/OCT scanning devices are disclosed in U.S. patentapplication Ser. No. 12/010,207 filed Jan. 22, 2008, now U.S. Pat. No.8,712,506, entitled MEDICAL IMAGING DEVICE WITH ROTARY ENCODER, which isincorporated herein by reference in its entirety.

Briefly, referring to FIGS. 26 a to 26 c, the imaging probes mayincorporate an encoder which is designed be used with an elongateimaging probe that uses a rotary shaft such as the imaging conduit 34 aspart of its scanning mechanism, its use can be generalized for use withany device that makes use of a long, flexible cable used fortransmission of torque where non-uniform rotational distortion may occurand an accurate estimation of rotary motion is required. In particular,it is most suited for use with flexible torque transmission systemswhere the outer diameter of the torque cable is relatively small (e.g.less than 4 mm) and long (e.g. longer than 5 cm) such that conventionalrotary encoding systems would not provide the desired angular resolutionor be adequately compact for the intended use.

FIG. 26 a demonstrates a longitudinal cross-section of the proximal anddistal ends of an elongate imaging device 450 with a torque transmissionshaft 451, mechanically coupled to a torque source 452. The torquesource 452 can be a motor, a handle that is manually turned by theoperator or any other such device. The torque transmission shaft 452transmits torque to the functional end 454 of the device, which can bean energy delivery device, a needle, an atherectomy head or any ofseveral other implements. In FIG. 26 c, the wall of an external sheath453 is shown to surround the transmission shaft and is shown to enclosethe functional end of the device although embodiments where the externalsheath is open or has openings near the functional end are possible. Anoptical fiber 455 is shown to be included as part of the external sheath453 for the purposes of enabling either the emitting light, detectinglight or both to travel to or from the encoding interface 104 that isremote to the proximal end of the transmission sheath. In FIG. 26 a thecylindrical encoding interface body 180 in this case is attached to arotating portion of the device while the fiber is relatively stationary.The optical fiber 455 may be included as part of the extrusion of theexternal sheath 453, as shown, or may be added to the inner or outersurface of the sheath and anchored to the sheath 453 by methods wellknown in the art, such as bonding or surrounding the fiber and sheathwith an additional layer of heat shrinkable material. The optical fiber455 is terminated with any necessary distal optics 115, such as anoptical spacer, lens and/or deflecting mechanism 172 (such as a prism ormirror) to direct light towards the encoding interface 104. The encodinginterface 104 in FIG. 26 a may be similar to that on the cylindricalencoding interface body disclosed in U.S. patent application Ser. No.12/010,207 filed Jan. 22, 2008, now U.S. Pat. No. 8,712,506, entitledMEDICAL IMAGING DEVICE WITH ROTARY ENCODER, mentioned above.

The encoding interface 104 in FIG. 26 b is similar to that on thecylindrical encoding interface body in the above mentioned copendingapplication. As the encoding optical circuit used in the embodiments ofFIGS. 14 a and 14 b are not mounted onto or directly coupled with thetorque transmission shaft, there is no need for an optical rotary jointalong the optical encoding circuit.

FIG. 26 c shows a cross-sectional image of a representativecross-section through the device 450 in FIG. 26 b through line 14 c-14c. One or more fiber optics 455 for the encoding system may beincorporated with the external sheath 453.

Thus the rotary encoder embodiments disclosed in U.S. patent applicationSer. No. 12/010,207 filed Jan. 22, 2008, now U.S. Pat. No. 8,712,506,entitled MEDICAL IMAGING DEVICE WITH ROTARY ENCODER, mentioned above canbe incorporated into an imaging probe 12 by substituting the functionalend of any of the embodiments in FIGS. 26 a to 26 d for an imagingassembly 30 and substituting the torque transmission shaft 451 for animaging conduit 34 suitable for carrying either electrical or opticalsignals.

As used herein, the terms “comprises”, “comprising”, “includes” and“including” are to be construed as being inclusive and open ended, andnot exclusive. Specifically, when used in this specification includingclaims, the terms “comprises”, “comprising”, “includes” and “including”and variations thereof mean the specified features, steps or componentsare included. These terms are not to be interpreted to exclude thepresence of other features, steps or components.

The foregoing description of the preferred embodiments of the inventionhas been presented to illustrate the principles of the invention and notto limit the invention to the particular embodiment illustrated. It isintended that the scope of the invention be defined by all of theembodiments encompassed within the following claims and theirequivalents.

What is therefore claimed is:
 1. A method of displaying co-registeredimages, said method comprising the steps of: obtaining a first image anda second image from an imaging catheter configured to obtain imagesaccording to two or more imaging modalities, wherein said first image isobtained according to a first imaging modality and said second image isobtained according to a second imaging modality, and wherein said firstimage and said second image are spatially co-registered; and dynamicallydisplaying an image comprising one or more portions of said first imageand one or more portions of said second image.
 2. The method accordingto claim 1 wherein said one or more portions of said first image andsaid one or more portions of said second image vary with time.
 3. Themethod according to claim 2 wherein said step of dynamically displayingsaid image comprises the steps of: dividing said image into a pluralityof sectors, wherein alternating sectors are displayed using alternatingimaging modalities; and displaying said image while varying locations ofsaid sectors.
 4. The method according to claim 3 wherein said sectorsrotate over time.
 5. The method according to claim 1 wherein said one ormore portions of said first image and said one or more portions of saidsecond image are determined according to input from a user.
 6. Themethod according to claim 5 wherein said user input comprises anidentification of one or more contours.
 7. The method according to claim1 wherein said step of dynamically displaying said image comprises thesteps of varying a transparency level of one or more of said first imageand said second image.
 8. The method according to claim 1 furthercomprising the steps of: processing one or more of said first image andsaid second image to identify one or more tissue types; generating anupdated image comprising an indication of said one or more tissue types;and displaying said updated image.
 9. A method of displaying a firstimage based on a feature that is present in a co-registered secondimage, said method comprising the steps of: obtaining said first imageand said second image from an imaging catheter configured to obtainimages according to two or more imaging modalities, wherein said firstimage is obtained according to a first imaging modality and said secondimage is obtained according to a second imaging modality, and whereinsaid first image and said second image are spatially co-registered;identifying said feature in said second image; identifying a location ofsaid feature in said first image; generating an image based on saidfirst image, said image comprising a superimposed identification of saidfeature; and displaying said image.
 10. The method according to claim 9wherein said second image comprises three-dimensional image data, andwherein said feature comprises one of a surface and a volume.
 11. Themethod according to claim 9 wherein said step of identifying saidfeature in said second image comprises receiving input from a user, saidinput identifying a location of said feature.
 12. The method accordingto claim 9 wherein said step of identifying said feature in said secondimage comprises a semi-automated identification of said feature, whereinsaid second image is processed to identify said feature with additionalinput from a user.
 13. The method according to claim 12 wherein saidadditional input from said user comprises one of a starting point for animage processing algorithm and a refinement of results produced by animage processing algorithm.
 14. The method according to claim 9 whereinsaid step of identifying said feature in said second image comprisesprocessing said second image to automatically identify said feature. 15.The method according to claim 14 wherein said step of processing saidsecond image comprises determining a location of a border in said secondimage.
 16. The method according to claim 14 wherein said step ofprocessing said second image comprises performing a pattern recognitionalgorithm to identify said feature in said second image.
 17. The methodaccording to claim 14 wherein said step of processing said second imagecomprises identifying one or more tissue types in said second image. 18.The method according to claim 14 wherein said second imaging modality isan acoustic imaging modality, and wherein said feature is identifiedaccording to a frequency-domain parameter.
 19. The method according toclaim 18 wherein said parameter is selected from the group consisting ofmaximum power, frequency of maximum power, minimum power, frequency ofminimum power, slope, y-intercept, mid-band fit and integratedback-scatter.
 20. The method according to claim 14 wherein said secondimaging modality is an optical imaging modality, and wherein saidfeature is identified according to an optical parameter.
 21. The methodaccording to claim 20 wherein said optical imaging modality is opticalcoherence tomography and wherein said first imaging modality isintravascular ultrasound.
 22. The method according to claim 21 whereinsaid optical parameter is selected from the group consisting of opticalcoherence tomography attenuation and optical coherence tomographypolarization sensitivity.
 23. The method according to claim 20 whereinsaid optical parameter is selected from the group consisting ofattenuation, polarization sensitivity, and spectroscopic information.24. The method according to claim 14 wherein said second imagingmodality is an optical imaging modality, and wherein said feature isidentified according to an edge detection algorithm.
 25. A method ofobtaining co-registered image data, said method comprising the steps of:obtaining a first image and a second image from an imaging catheterconfigured to obtain images according to two or more imaging modalities,wherein said first image and said second image are spatiallyco-registered; processing said first image and said second image toidentify one or more features in said first image and said second image;and spatial morphing one or more of said first image and said secondimage according to locations of said one or more features in said firstimage and said second image.
 26. A method of displaying features presentin first image and co-registered second image, said method comprisingthe steps of: a) obtaining said first image and said second image froman imaging catheter configured to obtain images according to two or moreimaging modalities, wherein said first image is obtained according to afirst imaging modality and said second image is obtained according to asecond imaging modality, and wherein said first image and said secondimage are spatially co-registered; b) processing said first image andsaid second image to identify a feature in said first image and saidsecond image; c) generating an image comprising said feature; and d)displaying said image.
 27. The method according to claim 26 wherein saidstep of processing said first image and said second image comprisesperforming a pattern recognition algorithm on said first image and saidsecond image to identify said feature.
 28. The method according to claim27 wherein said step of performing a pattern recognition algorithmfurther comprises determining a tissue type of said feature.
 29. Themethod according to claim 26 wherein said first imaging modality isoptical coherence tomography and said second imaging modality isintravascular ultrasound.